Radiography flat panel detector having a low weight X-ray shield and the method
of production thereof.
Technical Field
[0001] The present invention relates to diagnostic imaging and more particularly,
to a radiography X-ray detector having an X-ray shield which protects the
detector electronics and reduces or eliminates the impact of backscattered
X-rays during the exposure of the subject to the X-ray source.
Background Art
[0002] X-ray imaging is a non-invasive technique to capture medical images of
patients or animals as well as to inspect the contents of sealed containers,
such as luggage, packages, and other parcels. To capture these images,
an X-ray beam irradiates an object. The X-rays are then attenuated as
they pass through the object. The degree of attenuation varies across the
object as a result of variances in the internal composition and/or thickness
of the object. The attenuated X-ray beam impinges upon an X-ray detector designed to convert the attenuated beam to a usable shadow image of the
internal structure of the object.
[0003] Increasingly, radiography flat panel detectors (RFPDs) are being used to
capture images of objects during inspection procedures or of body parts of
patients to be analyzed. These detectors can convert the X-rays directly
into electric charges (direct conversion direct radiography - DCDR), or in
an indirect way (indirect conversion direct radiography - ICDR).
[0004] In direct conversion direct radiography, the RFPDs convert X-rays directly
into electric charges. The X-rays are directly interacting with a
photoconductive layer such as amorphous selenium (a-Se).
[0005] In indirect conversion direct radiography, the RFPDs have a scintillating
phosphor such as Csl:TI (caesium iodide doped with thallium) or Gd20 2S
(gadolinium oxysulphide) which converts X-rays into light which then
interacts with an amorphous silicon (a-Si) semiconductor layer, where
electric charges are created.
[0006] The created electric charges are collected via a switching array,
comprising thin film transistors (TFTs). The transistors are switched-on
row by row and column by column to read out the signal of the detector.
The charges are transformed into voltage, which is converted in a digital
number that is stored in a computer file which can be used to generate a
softcopy or hardcopy image. Recently Complementary Metal Oxides
Semiconductors (CMOS) sensors are becoming important in X-ray
imaging. The detectors based on CMOS are already used in
mammography, dental, fluoroscopy, cardiology and angiography images.
The advantage of using those detectors is a high readout speed and a low
electronic noise.
[0007] Generally, the imaging array including TFTs as switching array and
photodiodes (in case of ICDR) is deposited on a thin substrate of glass.
The assembly of scintillator or photoconductor and the imaging array on
the glass substrate does not absorb all primary radiation, coming from the
X-ray source and transmitted by the object of the diagnosis. Hence the
electronics positioned under this assembly are exposed to a certain
fraction of the primary X-ray radiation. Since the electronics are not
sufficiently radiation hard, this transmitted radiation may cause damage.
[0008] Moreover, X-rays which are not absorbed by the assembly of scintillator or
photoconductor and the imaging array on the glass substrate, can be
absorbed in the structures underneath the glass substrate. The primary
radiation absorbed in these structures generates secondary radiation that
is emitted isotropically and that thus exposes the imaging part of the
detector. The secondary radiation is called "backscatter" and can expose
the image part of the detector thereby introducing artefacts into the
reconstructed image. Since the space under the assembly is not
homogeneously filled, the amount of scattered radiation is position
dependent. Part of the scattered radiation is emitted in the direction of the
assembly of scintillator or photoconductor and imaging array and may
contribute to the recorded signal. Since this contribution is not spatially
homogeneous this contribution will lead to haze in the image, and,
therefore, reduce the dynamic range. It will also create image artefacts.
[0009] To avoid damage to the electronics and image artefacts due to scattered
radiation, an X-ray shield may be applied underneath the assembly of
scintillator or photoconductor and imaging array. Because of their high
density and high intrinsic stopping power for X-rays, metals with a high
atomic number are used as materials in such an X-ray shield. Examples of
these are sheets or plates from tantalum, lead or tungsten as disclosed in
EP1471384B1 , US2013/0032724A1 , US2012/0097857A1 .
[0010] However, metals with a high atomic number also have a high density.
Hence, X-ray shields based on these materials have a high weight. Weight
is an important characteristic of the RFPD especially for the portability of
the RFPDs. Any weight reduction is, therefore, beneficial for the users of
the RFPDs such as medical staff.
[001 1] US7317190B2 discloses a radiation absorbing X-ray detector panel
support comprising a radiation absorbing material to reduce the reflection
of X-rays of the back cover of the X-ray detector. The absorbing material
including heavy atoms such as lead, barium sulphate and tungsten can be
disposed as a film via a chemical vapour deposition technique onto a rigid
panel support or can be mixed via injection moulding with the base
materials used to fabricate the rigid panel support. The support for the
chemical vapour deposition as well as the base materials to fabricate the
rigid panel support, represent an extra weight contribution in the RFPD.
Moreover, the detector panel support comprising the radiation absorbing
material needs to be additionally fixed to assure immobilisation to the
detector.
[0012] In US5650626, an X-ray imaging detector is disclosed which contains a
substrate, supporting the conversion and detection unit. The substrate
includes one or more elements having atomic numbers greater than 22.
Since the detection array is directly deposited on the substrate, the variety
of suitable materials of the substrate is rather limited.
[0013] In US5777335, an imaging device is disclosed comprising a substrate,
preferably glass containing a metal selected from a group formed by Pb,
Ba, Ta or W. According to the inventors, the use of this glass would not
require an additional X-ray shield based on lead. However, glass
containing sufficient amounts of metals from a group formed by Pb, Ba, Ta
or is more expensive than glass which is normally used as a substrate
for imaging arrays.
[0014] US7569832 discloses a radiographic imaging device, namely a RFPD,
comprising two scintillating phosphor layers as scintillators each one
having different thicknesses and a transparent substrate to the X-rays
between said two layers. The use of an additional phosphor layer at the
opposite side of the substrate improves the X-ray absorption while
maintaining the spatial resolution. The presence of the additional phosphor
layer as disclosed is not sufficient to absorb all primary X-ray radiation to
prevent damage of the underlying electronics and to prevent backscatter.
An extra X-ray shield will still be required in the design of this RFPD.
[0015] In US2008/01 1960A1 a dual-screen digital radiography apparatus is
claimed. This apparatus consists of two flat panel detectors (front panel
and back panel) each comprising a scintillating phosphor layer to capture
and process X-rays. The scintillating phosphor layer in the back panel
contributes to the image formation and has no function as X-ray shield to
protect the underlying electronics. This dual-screen digital flat panel, still
requires an X-ray shield to protect the underlying electronics and to avoid
image artefacts due to scattered radiation.
[001 6] WO20051 055938 discloses a light weight film, with an X-ray absorption at
least equivalent to 0.254 mm of lead and which has to be applied on
garments or fabrics for personal radiation protection or attenuation, such
as aprons, thyroid shields, gonad shields, gloves, etc. Said film is obtained
from a polymer latex mixture comprising high atomic weight metals or their
related compounds and/or alloys. The suitable metals are the ones that
have an atomic number greater than 45. No use of this light weight film in
a RFPD is mentioned. Although a light weight film is claimed, the metal
particles used in the composition of the film still contribute to a high extend
to the weight of the shield.
[00 7] US6548570 discloses a radiation shielding composition to be applied on
garments or fabrics for personal radiation protection. The composition
comprised a polymer, preferably an elastomer, and a homogeneously
dispersed powder of a metal with high atomic number in an amount of at
least 80% in weight of the composition as filler. A loading material is mixed
with the filler material and kneaded with the elastomer at a temperature
below 180°C resulting in a radiation shielding composition that can be
applied homogeneously to garments and fabrics on an industrial scale.
The use of metals is however increasing the weight of the shield of this
invention considerably.
[0018] WO2009/0078891 discloses a radiation shielding sheet which is free from
lead and other harmful components having a highly radiation shielding
performance and an excellent economical efficiency. Said sheet is formed
by filling a shielding material into an organic polymer material, the
shielding material being an oxide powder containing at least one element
selected from the group consisting of lanthanum (La), cerium (Ce),
praseodymium (Pr), neodymium (Nd), samarium (Sm), europium (Eu) and
gadolinium (Gd) and the polymer being a material such as rubber,
thermoplastic elastomer, polymer resin or similar. The volumetric amount
of the shielding material filled in the radiation shielding sheet is 40 to 80
vol. % with respect to the total volume of the sheet. No use of this film into
a RFPD is mentioned.
[0019] From the foregoing discussion, it should be apparent that there is a need
for a RFPD with an X-ray shield to protect the underlying electronics and
to absorb the scattered radiation produced by the underlying structures to
avoid image artefacts in the imaging area, but which has a low weight, a
low cost, which can be produced in an economically efficient way and
which does not have to be fixed to the substrate of the imaging array in an
additional step of the production.
Summary of invention
[0020] It is therefore an object of the present invention to provide a solution for
the high weight contribution of the X-ray shield in a radiography flat panel
detector having a single imaging array and to provide at the same time a
solution for producing the RFPD on an economically efficient way. The
object has been achieved by a radiography flat panel detector as defined
in claim 1.
[0021] An additional advantage of the RFPD as defined in claiml , is that the
thickness of said X-ray shield can be adjusted in a continuous way to the
required degree of the X-ray shielding effect instead of in large steps as it
is in the case of shielding metal sheets commercially available with
standard thicknesses. Even though plates with custom made thickness
can be purchased, the price of those metal plates is still very high because
of the customization.
[0022] According to another aspect, the present invention includes a method of
manufacturing a radiography flat panel detector. The method includes
coating or depositing on the substrate of the imaging array, preferably on
the opposite side of the imaging array, an X-ray absorbing layer with at
least one chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements and which has
a dimensionless absorption exponent for 60 keV Am241 source greater
than 0.5 as defined in claim 1.
[0023] Other features, elements, steps, characteristics and advantages of the
present invention will become more apparent from the following detailed
description of preferred embodiments of the present invention. Specific
embodiments of the invention are also defined in the dependent claims.
Brief description of drawings
[0024] Fig. 1 represents a cross-section of a RFPD according to one embodiment
of the present invention and the underlying electronics, wherein:
1 is a scintillator or photoconductive layer
2 is single imaging array
3 is a substrate
4 is an X-ray absorbing layer
5 is the underlying electronics
Description of embodiments
[0025] The present invention relates to a radiography flat panel detector (RFPD)
comprising a scintillator or photoconductive layer, a single imaging array
on a substrate and an X-ray shield having an X-ray absorbing layer
comprising a chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements coated or
deposited on a side of a substrate of an imaging array. If the chemical
compound in the X-ray absorbing layer is a scintillating phosphor, a layer
is present between the X-ray absorbing layer and the substrate, which has
a transmission for light of 10% or lower at the wavelength of the light
emission of said chemical compound.
The X-ray absorbing layer
[0026] It has been found that X-ray shields can be made with the same X-ray
stopping power but with considerably less weight than X-ray shields
consisting of metals only by use of a layer comprising one or more
chemical compounds having a metal element with an atomic number of 20
or more and one or more non-metal elements. Preferably these
compounds are oxides or salts such as halides, oxysulphides, sulphites,
carbonates of metals with an atomic number of 20 or higher. Examples of
suitable metal elements with an atomic number higher than 20 that can be
used in the scope of the present invention are metals such as Barium (Ba),
Calcium (Ca), Cerium (Ce), Caesium (Cs), Gadolinium (Gd), Lanthanum
(La), Lutetium (Lu), Palladium (Pd), Tin (Sn), Strontium (Sr), Tellurium
(Te), Yttrium (Y), and Zinc (Zn). A further advantage of the invention is that
these compounds are relatively inexpensive and are characterised by a
low toxicity.
[0027] Examples of preferred compounds having a metal element with an atomic
number of 20 or more and one or more non-metal elements, are Caesium
iodide (Csl), Gadolinium oxysulphide (Gd202S), Barium fluorobromide
(BaFBr), Calcium tungstate (CaWCU), Barium titanate (BaTi0 3) ,
Gadolinium oxide (Gd203), Barium chloride (BaCb), Barium fluoride
(BaF2), Barium oxide (BaO), Cerium oxides, Caesium nitrate (CSNO3),
Gadolinium fluoride (GdF2) , Palladium iodide (Pdl2) , Tellurium dioxide
(Te0 2) , Tin iodides, Tin oxides, Barium sulphides, Barium carbonate
(BaC03), Barium iodide, Caesium chloride (CsCI), Caesium bromide
(CsBr), Caesium fluoride (CsF), Caesium sulphate (Cs2S0 4) , Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides, Rhenium sulphides, BaFX (wherein X represents CI or I), RFXn
(wherein RF represents lanthanides selected from: La, Ce, Pr, Nd, Pm,
Sm, Eu, Gd, Tb, Dy, Ho, Er, Tm, Yb, Lu and X represents halides selected
from: F, CI, Br, I), RFyOz, RFy(SO4)z, RFySz and/or RFy(WO4)z, wherein n,
y, z are independently an integer number higher than 1. These
compounds can produce lower weight X-ray shields and are easy to
handle due to their low hygroscopicity than their pure metal analogues.
The most preferred metallic compounds are: Gd2O2S, Gd2O3, Ce2O3, Csl,
BaFBr, CaWO and BaO.
[0028] It is another advantage of the present invention that the range of metal
elements which can be used for the x-ray absorbing layer, is much larger
than the corresponding range of the pure metals and/or alloys, since many
of them are not stable in their elemental form. Examples are the alkali
metals, the alkaline earth metals and the rare-earth metals.
[0029] The chemical compounds having a metal element with an atomic number
of 20 or more and one or more non-metal elements may be used in the Xray
absorbing layer of the present invention as powder dispersed in a
binder. The amount of the binder in the X-ray absorbing layer in weight
percent can vary in the range from 1% to 50%, preferably from 1% to 25%,
more preferably from 1% to 10%, most preferably from 1% to 3%.
[0030] Suitable binders are e.g. organic polymers or inorganic binding
components. Examples of suitable organic polymers are polyethylene
glycol acrylate, acrylic acid, butenoic acid, propenoic acid, urethane
acrylate, hexanediol diacrylate, copolyester tetracrylate, methylated
melamine, ethyl acetate, methyl methacrylate. Inorganic binding
components may be used as well. Examples of suitable inorganic binding
components are alumina, silica or alumina nanoparticles, aluminium
phosphate, sodium borate, barium phosphate, phosphoric acid, barium
nitrate.
[0031] Preferred binders are organic polymers such as cellulose acetate butyrate,
polyalkyl (meth)acrylates, polyvinyl-n-butyral, poly(vinylacetate-covinylchloride),
poly(acrylonitrile-co-butadiene-co-styrene), polyvinyl
chloride-co-vinyl acetate-co-vinylalcohol), poly(butyl acrylate), poly(ethyl
acrylate), poly(methacrylic acid), polyvinyl butyral), trimellitic acid,
butenedioic anhydride, phtalic anhydride, polyisoprene and/or a mixture
thereof. Preferably, the binder comprises one or more styrenehydrogenated
diene block copolymers, having a saturated rubber block
from polybutadiene or polyisoprene, as rubbery and/or elastomeric
polymers. Particularly suitable thermoplastic rubbers, which can be used
as block-copolymeric binders, in accordance with this invention, are the
KRATON™ G rubbers, KRATON™ being a trade name from SHELL.
[0032] In case the coating of the X-ray absorbing layer is to be cured, the binder
includes preferably a polymerisable compound which can be a
monofunctional or polyfunctional monomer, oligomer or polymer or a
combination thereof. The polymerisable compounds may comprise one or
more polymerisable groups, preferably radically polymerisable groups. Any
polymerisable mono- or oligofunctional monomer or oligomer commonly
known in the art may be employed. Preferred monofunctional monomers
are described in EP1637322A paragraph [0054] to [0057]. Preferred
oligofunctional monomers or oligomers are described in EP1637322A
paragraphs [0059] to [0064]. Particularly preferred polymerisable
compound are urethane (meth)acrylates and ,6-hexanedioldiacrylate.
The urethane (meth)acrylates are oligomer which may have one, two,
three or more polymerisable groups.
[0033] Suitable solvents, to dissolve the binder being an organic polymer during
the preparation of the coating solution of the X-ray absorbing layer can be
acetone, hexane, methyl acetate, ethyl acetate, isopropanol, methoxy
propanol, isobutyl acetate, ethanol, methanol, methylene chloride and
water. The most preferable ones are toluene, methyl-ethyl-ketone (MEK)
and methyl cyclohexane. To dissolve suitable inorganic binding
components, water is preferable as the main solvent. In case of a curable
coating liquid, one or more mono and/or difunctional monomers and/or
oligomers can be used as diluents. Preferred monomers and/or oligomers
acting as diluents are miscible with the above described urethane
(meth)acrylate oligomers. The monomer(s) or oligomer(s) used as diluents
are preferably low viscosity acrylate monomer(s).
[0034] The X-ray absorbing layer of the present invention may also comprise
additional compounds such as dispersants, plasticizers, photoinitiators,
photocurable monomers, antistatic agents, surfactants, stabilizers
oxidizing agents, adhesive agents, blocking agents and/or elastomers.
[0035] Dispersants which can be used in the present invention include nonsurface
active polymers or surface-active substances such as surfactants,
added to the binder to improve the separation of the particles of the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements and to further prevent
settling or clumping in the coating solution. Suitable examples of
dispersants are Stann JF95B from Sakyo and Disperse Ayd™ 1900 from
Daniel Produkts Company. The addition of dispersants to the coating
solution of the X-ray absorbing layer improves further the homogeneity of
the layer.
[0036] Suitable examples of plasticizers are Plastilit™ 3060 from BASF,
Santicizer™ 278 from Solutia Europe and Palatinol™ C from BASF. The
presence of plasticizers to the X-ray absorbing layer improves the
compatibility with flexible substrates.
[0037] Suitable photo-initiators are disclosed in e.g. J.V. Crivello et al. in "
Photoinitiators for Free Radical, Cationic & Anionic Photopolymerisation
2nd edition", Volume III of the Wiley/SITA Series In Surface Coatings
Technology, edited by G. Bradley and published in 1998 by John Wiley
and Sons Ltd London, pages 276 to 294.Examples of suitable
photoinitiators can be Darocure™ 1173 and Nuvopol™ PI-3000 from
Rahn. Examples of suitable antistatic agents can be Cyastat™ SN50 from
Acris and Lanco™ STAT K 100N from Langer.
[0038] Examples of suitable surfactants can be Dow Corning™ 190 and Gafac
RM710, Rhodafac™ RS-710 from Rodia. Examples of suitable stabilizer
compounds can be Brij™ 72 from ICI Surfactants and Barostab™ MS from
Baerlocher Italia. An example of a suitable oxidizing agent can be lead (IV)
oxide from Riedel De Haen. Examples of suitable adhesive agents can be
Craynor™ 435 from Cray Valley and Lanco™ wax TF1780 from Noveon.
An example of a suitable blocking agent can be Trixene™ BI7951 from
Baxenden. An example of a suitable elastomer compound can be Metaline
™from Schramm).
The thickness of the X-ray absorbing layer, the atomic number of the metal
element and the concentration of the chemical compound having a metal
element with an atomic number of 20 or more can be chosen to achieve a
desired level of X-ray absorption or attenuation in the RFPD. The value of
this level can be expressed as the "absorption exponent" (AE) and should
be equal to or higher than 0.5 to protect sufficiently the underlying
electronics of the RFPD and to limit the impact from backscattered X-rays
on the obtained image. The absorption exponent is a physical parameter
that is equal to the negative of the natural logarithm of the X-ray
transmittance. Since transmittance varies with X-ray energy, the
absorption exponent is more conveniently expressed relative to X-rays
emitted by a standard radiation source. A convenient standard is the 59.57
keV (hereafter 60 keV) gamma ray emission of Am241 . This source is in the
middle range of X-ray energies typically used in medical imaging, 20 to
150 keV, and is commonly used as a source of monoenergetic X-rays for
experiments. The absorption exponent can be measured directly or can be
calculated using formula 1 (expressed here for a 60 keV gamma ray
emission Am241 source):
AE( Am24 60keV)= t*(kiei +k2e2+k3e3+. ..) (Formula )
wherein AE(Am241 60 keV) represents the absorption exponent of the
substrate relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer in the principle
direction of propagation of the primary X-ray beam; e-i, b 2, 3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and
ki,k2,k3... represent the mass attenuation coefficients of the respective
elements at given energy. As the formula indicates, the absorption
exponent is equal to a thickness dimension multiplied by the sum of the
products of the mass attenuation coefficient for each element in the X-ray
absorbing layer at the about 60 keV gamma ray emission of Am241 and the
respective concentration of each element in the X-ray absorbing layer. The
absorption exponent is dimensionless. For example, if the mass
attenuation coefficients are expressed in cm2/mole, the concentrations
should be expressed in moles/cm3 and the thickness in centimetres. Mass
attenuation coefficients can be found on the 'National Institute for
Standards and Technology' (www.nist.gov/pml/data/xraycoef/). Depending
on the application, the coating weight of the chemical compound having a
metal element with an atomic number of 20 or more and one or more non
metal elements in the X-ray absorbing layer can be flexibly adjusted and in
case of using a RFPD for medical purposes, this coating weight is
preferably at least 100 mg/cm2, more preferably at least 200 mg/cm2.
[0040] The thickness of the X-ray absorbing layer can vary as well and depends
on the necessary shielding power and/or the space available to
incorporate the X-ray shield in the design of the RFPD. In the present
invention, the thickness of the X-ray absorbing layer can be at least 0.1
mm, more preferably in the range from 0.1 mm to 2.0 mm.
The light absorbing or light reflecting layer
[0041] Some of the chemical compounds having a metal element with an atomic
number of 20 or more and at least one non-metal elements are scintillating
phosphors which can emit light on X-ray absorption. If this is the case, light
emitted by these scintillating phosphors in the X-ray absorbing layer can
reach the imaging array through the substrate and contribute to the image
formation. Due to scattering in the substrate of the imaging array of the
light emitted by the scintillating phosphor present in the X-ray absorbing
layer, the quality of the image of the investigated object is negatively
impacted. In the case that scintillating phosphors are present in the X-ray
absorbing layer, a light reflecting or light absorbing layer is to be present
between the X-ray absorbing layer and the imaging array, more preferably
between the X-ray absorbing layer and the substrate of the imaging array.
In order to avoid any contribution of the emitted light by scintillating
phosphors in the X-ray absorbing layer to the image, the transmission of
the emitted light from the scintillating phosphor through this light absorbing
or reflecting layer, should be equal to or lower than 10%, more preferable
lower than 3%, most preferably lower than 1%. The term 'scintillating
phosphor' in the X-ray absorbing layer according to the invention should
be interpreted as a compound whose light emission on X-ray absorption
can reach the imaging array and contribute to the image formation of the
detector.
[0042] White coloured layers may be used to reflect light emitted by the
scintillating phosphor in the X-ray absorbing layer. Layers comprising T1O2
are preferably used to reflect 90% or more light at the wavelength(s) of the
light emitted by the scintillating phosphor. The solid content of T1O2 in the
light reflecting layer is preferably in the range of 25 to 50 (wt.)%. and the
thickness is preferably in the range of 5 to 40 pm. More preferably, the
solid content of the T1O2 is 33 to 38(wt.)% of the total solid content of the
layer and the layer thickness is between 13 and 30 pm. The layer is
preferably applied with a doctor blade coater on the substrate of the
imaging array, preferably on the side opposite to the imaging array.
[0043] In another preferred embodiment of the invention, black coloured layers
can be used to absorb light emitted by a scintillating phosphor in the X-ray
absorbing layer because of their high efficiency to absorb light. Black
particles, such as fine carbon black powder (ivory black, titanium black,
iron black), are suitable to obtain sufficient absorption of emitted light by
the scintillating phosphor. Preferably the solid content of carbon black is in
the range of 3 to 30 (wt.)% and a layer thickness of 2 to 30 pm will absorb
90% or more of the emitted light by the scintillating phosphor. More
preferably the range of the solid content of the carbon black is in the range
of 6 to 15 (wt.)% and the layer thickness between 5 and15 pm. In another
embodiment of the invention, coloured pigments or dyes absorbing
specifically at the maximal wavelength of the emitted light by the
scintillating phosphor in the X-ray absorbing layer can be used.
The scintillator
[0044] In the RFPD for indirect conversion direct radiography according to the
present invention, the scintillator comprises optionally a support and
provided thereon, a scintillating phosphor such as one or more of
Gd20 2S:Tb, Gd20 2S:Eu, Gd203:Eu, La20 2S:Tb, La20 2S, Y20 2S:Tb,
Cs TI, Csl:Eu, Csl:Na, CsBr:TI, Nal:TI, CaW0 , CaW0 :Tb, BaFBnEu,
BaFCLEu, BaS0 4:Eu, BaSrS0 4, BaPbS0 4, BaAh20i9 :Mn,
BaMgAlioOi 7:Eu, Zn2Si0 :Mn, (Zn, Cd)S:Ag, LaOBr, LaOBr:Tm,
Lu2O2S:Eu, Lu20 2S:Tb, LuTa04, Hf0 2:Ti, HfGe0 :Ti, YTa04, YTa04:Gd,
YTa04:Nb, Y203:Eu, YB0 3:Eu, YB0 3:Tb, or (Y,Gd)B0 3:Eu, or
combinations thereof. Besides crystalline scintillating phosphors,
scintillating glass or organic scintillators can also be used.
[0045] When evaporated under appropriate conditions, a layer of doped Csl will
condense in the form of needle like, closely packed crystallites with high
packing density onto a support. Such a columnar or needle-like scintillating
phosphor is known in the art. See, for example, ALN Stevels et al. , "Vapor
Deposited Csl:Na Layers: Screens for Application in X-Ray Imaging
Devices, " Philips Research Reports 29:353-362 (1974); and T. Jing et al,
"Enhanced Columnar Structure in Csl Layer by Substrate Patterning",
IEEE Trans. Nucl. Sci. 39: 1 95- 198 (1992). More preferably, the
scintillating phosphor layer includes doped Csl.
[0046] A blend of different scintillating phosphors can also be used. The median
particle size is generally between about 0. 5 m ti and about 40 mih . A
median particle size of between 1 mi and about 20 mih is preferred for
ease of formulation, as well as optimizing properties, such as speed,
sharpness and noise. The scintillator for the embodiments of the present
invention can be prepared using conventional coating techniques whereby
the scintillating phosphor powder, for example Gd20 2S is mixed with a
solution of a binder materia! and coated by means of a blade coater onto a
substrate. The binder can be chosen from a variety of known organic
polymers that are transparent to X-rays, stimulating, and emitting light.
Binders commonly employed in the art include sodium osulfobenzaldehyde
acetal of poly(vinyl alcohol); chloro-sulfonated
poly(ethylene); a mixture of macromolecular bisphenol poly(carbonates)
and copolymers comprising bisphenol carbonates and poly(alkylene
oxides);aqueous ethanol soluble nylons; poly(alkyl acrylates and
methacrylates) and copolymers of poly(alkyl acrylates and methacrylates
with acrylic and methacrylic acid); polyvinyl butyral); and poly(urethane)
elastomers. Other preferable binders which can be used are described
above in the section of the X-ray absorbing layer. Any conventional ratio
phosphor to binder can be employed. Generally, the thinner scintillating
phosphor layers are, the sharper images are realized when a high weight
ratio of phosphor to binder is employed. Phosphor-to-binder ratios in the
range of about 70:30 to 99:1 by weight are preferable.
The photoconductive layer
[0047] In the RFPD for direct conversion direct radiography according to the
present invention, the photoconductive layer is usually amorphous
selenium, although other photoconductors such as Hgl2, PbO, Pbl2, TIBr,
CdTe and gadolinium compounds can be used. The photoconductive layer
is preferentially deposited on the imaging array via vapour deposition but
can also been coated using any suitable coating method.
The imaging array and its substrate
[0048] The single imaging array used in the invention for indirect conversion
direct radiography is based on an indirect conversion process which uses
several physical components to convert X-rays into light that is
subsequently converted into electrical charges. First component is a
scintillating phosphor which converts X-rays into light (photons). Light is
further guided towards an amorphous silicon photodiode layer which
converts light into electrons and electrical charges are created. The
charges are collected and stored by the storage capacitors. A thin-film
transistor (TFT) array adjacent to amorphous silicon read out the electrical
charges and an image is created. Examples of suitable image arrays are
disclosed in US5262649 and by Samei E. et al., "General guidelines for
purchasing and acceptance testing of PACS equipment", Radiographics,
24, 313-334 . Preferably, the imaging arrays as described in
US201 3/0048866, paragraph [90-125] and US201 3/221 230, paragraphs
[53-71] and [81-104] can be used.
[0049] The imaging array used in the invention for direct conversion direct
radiography is based on a direct conversion process of X-ray photons into
electric charges. In this array, an electric field is created between a top
electrode, situated on top of the photoconductor layer and the TFT
elements. As X-rays strike the photoconductor, the electric charges are
created and the electrical field causes to move them towards the TFT
elements where they are collected and stored by storage capacitors.
Examples of suitable image arrays are disclosed by Samei E. et al., "
General guidelines for purchasing and acceptance testing of PACS
equipment", Radiographics, 24, 313-334.
[0050] For both the direct and indirect conversion process, the charges must be
read out by readout electronics. Examples of readout electronics in which
the electrical charges produced and stored are read out row by row, are
disclosed by Samei E. et al., Advances in Digital Radiography. RSNA
Categorical Course in Diagnostic Radiology Physics (p. 49-61) Oak Brook,
III.
[0051] The substrate of the imaging array of the present invention is preferably
glass. However, imaging arrays fabricated on substrates made of plastics,
metal foils can also be used. The imaging array can be protected from
humidity and environmental factors by a layer of silicon nitride or polymer
based coatings such as fluoropolymers, polyimides, polyamides,
polyurethanes and epoxy resins. Also polymers based on B-staged
bisbenzocyclobutene-based (BCB) monomers can be used. Alternatively,
porous inorganic dielectrics with low dielectric constants can also be used.
The underlying electronics
[0052] The underlying electronics, situated under the X-ray absorbing layer
comprise a circuit board which is equipped with electronic components for
processing the electrical signal from the imaging array, and/or controlling
the driver of the imaging array and is electrically connected to the imaging
array.
Method of making the radiographic flat panel detector
Method of making the X-ray shield
[0053] The X-ray shield of the present invention can be obtained by applying an
X-ray absorbing layer comprising at least one chemical compound having
a metal element with an atomic number of 20 or more and one or more
non-metal elements onto the substrate carrying the single imaging array.
Preferably, the X-ray absorbing layer is applied on the side of the
substrate opposite to the imaging array. Any known method for applying
layers on a substrate can be suitable, e.g. Physical Vapour Deposition
(PVD), Chemical Vapour Deposition (CVD), sputtering, doctor blade
coating, spin-coating, dip-coating, spray-coating, knife coating, screen
printing and lamination. The most preferable methods are doctor blade
coating and PVD.
[0054] One of the preferred methods of applying a layer is by coating a solution ,
hereafter denoted as coating solution, comprising the chemical compound
having a metal element with an atomic number of 20 or more and one or
more non-metal elements and a binder onto the substrate of the single
imaging array. In a preferred embodiment the coating solution is prepared
by first dissolving the binder in a suitable solvent. To this solution the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements is added. To obtain a
homogenous coating solution, a homogenization step or milling step of the
mixture can be included in the preparation process. A dispersant can be
added to the binder solution prior to the mixing with the chemical
compound having a metal element with an atomic number of 20 or more
and one or more non-metal elements. The dispersant improves the
separation of the particles in the coating solution and prevents settling or
clumping of the ingredients in the coating solution. The addition of
dispersants to the coating solution of the X-ray absorbing layer further
decreases the surface tension of the coating solution and improves the
coating quality of the X-ray absorbing layer.
[0055] In another embodiment of the invention, the binder being a polymerisable
compound can be dissolved in diluents comprising one or more mono
and/or difunctional monomers and/or oligomers.
[0056] After stirring or homogenization the coating solution is applied onto the
substrate preferably using a coating knife or a doctor blade. By adjusting
the distance between the coating blade and the substrate. After the
coating of the X-ray absorbing layer, this layer can be dried via an IRsource,
an UV-source, a heated metal roller or heated air. When
photocurable monomers are used in the coating solution, the coated layer
can be cured via heating or via an UV-source.
[0057] In another preferred embodiment, a PVD process is used in which the Xray
absorbing layer comprising the chemical compound having a metal
element with an atomic number of 20 or more and one or more non-metal
elements is prepared in vacuum from the gas phase of melting materials.
The material in a solid form can be introduced in a heat resistive container
to a vacuum chamber and subsequently heated to the temperature equal
to or higher than the melting point of compound(s).

Claims
. A radiography flat panel detector comprising a layer configuration in the order
given,
a) a scintillator or photoconductive layer (1)
b) a single imaging array (2)
c) a substrate (3)
d) an X-ray absorbing layer (4) comprising a chemical compound having a
metal element with an atomic number of 20 or more and one or more nonmetal
elements,
characterised in that the X-ray absorbing layer has a dimensionless absorption
exponent of greater than 0.5 for gamma ray emission of Am241 at about 60keV;
wherein
AE( A m 24 60keV)= t* (kiei+k 2 e 2+k3 e3+ . . . )
wherein AE(Am241 60 keV) represents the absorption exponent of the X-ray
absorbing layer relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer; e-i , e2, e3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and ki,k 2 ,k3. . .
represent the mass attenuation coefficients of the respective elements, and if
the chemical compound is a scintillating phosphor, a layer is present between
the X-ray absorbing layer and the substrate, the layer having a transmission for
light of 10% or lower at the wavelength of the light emission of the chemical
compound.
2 . The radiography flat panel detector according to claim 1, wherein the X-ray
absorbing layer (4) is positioned between the substrate (3) and the underlying
electronics (5).
3. The radiography flat panel detector according to claim 1 or 2, wherein the
chemical compound is selected form the group consisting of Csl, Gd20 2S,
BaFBr, CaW0 , BaTiO3, Gd20 3, BaCI2, BaF2, BaO, Ce20 3, Ce0 2, CsN0 3
GdF2, Pdl2, Te0 2, Snl2, SnO, BaS0 , BaCOs, Bal, BaFX, RFXn, RFyO ,
RFy(S04)z, RFySz, RFy(W0 )z, CsBr, CsCI, CsF, CsN0 3, Cs2S0 Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides and Rhenium sulphides or mixtures thereof, wherein:
X is a halide selected from the group of F, CI, Br and I ; and
RF is a lanthanide selected from La, Ce, Pr, Nd, Pm, Sm, Eu, Gd,
Tb, Dy, Ho, Er, Tm, Yb and Lu; and
n, y, z are independently an integer number higher than .
4. The radiography flat panel detector according to any of the preceding claims,
wherein the X-ray absorbing layer comprises a binder.
5. The radiography flat panel detector according to claim 4, wherein the amount
of the binder in the X-ray absorbing layer is 10% by weight or less.
6 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, comprises a dye
or a pigment.
7. The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, is light absorbing.
8 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light emitted by the chemical
compound of 10% or lower, comprises light reflecting particles.
9 . A method of making a radiography flat panel detector as defined in claim 1,
comprising the steps of:
a) providing a substrate (3) with an imaging array (2) on a side of the first
substrate; and
b) applying a scintillating phosphor (1) onto the imaging array; and
c) applying the X-ray absorbing layer (4) on the side of the substrate opposite
to the imaging array.
10. The method of making a radiography flat panel detector according to claim 9
wherein the X-ray absorbing layer is coated by means of knife coating or
doctor blade coating.designed to convert the attenuated beam to a usable shadow image of the
internal structure of the object.
[0003] Increasingly, radiography flat panel detectors (RFPDs) are being used to
capture images of objects during inspection procedures or of body parts of
patients to be analyzed. These detectors can convert the X-rays directly
into electric charges (direct conversion direct radiography - DCDR), or in
an indirect way (indirect conversion direct radiography - ICDR).
[0004] In direct conversion direct radiography, the RFPDs convert X-rays directly
into electric charges. The X-rays are directly interacting with a
photoconductive layer such as amorphous selenium (a-Se).
[0005] In indirect conversion direct radiography, the RFPDs have a scintillating
phosphor such as Csl:TI (caesium iodide doped with thallium) or Gd20 2S
(gadolinium oxysulphide) which converts X-rays into light which then
interacts with an amorphous silicon (a-Si) semiconductor layer, where
electric charges are created.
[0006] The created electric charges are collected via a switching array,
comprising thin film transistors (TFTs). The transistors are switched-on
row by row and column by column to read out the signal of the detector.
The charges are transformed into voltage, which is converted in a digital
number that is stored in a computer file which can be used to generate a
softcopy or hardcopy image. Recently Complementary Metal Oxides
Semiconductors (CMOS) sensors are becoming important in X-ray
imaging. The detectors based on CMOS are already used in
mammography, dental, fluoroscopy, cardiology and angiography images.
The advantage of using those detectors is a high readout speed and a low
electronic noise.
[0007] Generally, the imaging array including TFTs as switching array and
photodiodes (in case of ICDR) is deposited on a thin substrate of glass.
The assembly of scintillator or photoconductor and the imaging array on
the glass substrate does not absorb all primary radiation, coming from the
X-ray source and transmitted by the object of the diagnosis. Hence the
electronics positioned under this assembly are exposed to a certain
fraction of the primary X-ray radiation. Since the electronics are not
sufficiently radiation hard, this transmitted radiation may cause damage.
[0008] Moreover, X-rays which are not absorbed by the assembly of scintillator or
photoconductor and the imaging array on the glass substrate, can be
absorbed in the structures underneath the glass substrate. The primary
radiation absorbed in these structures generates secondary radiation that
is emitted isotropically and that thus exposes the imaging part of the
detector. The secondary radiation is called "backscatter" and can expose
the image part of the detector thereby introducing artefacts into the
reconstructed image. Since the space under the assembly is not
homogeneously filled, the amount of scattered radiation is position
dependent. Part of the scattered radiation is emitted in the direction of the
assembly of scintillator or photoconductor and imaging array and may
contribute to the recorded signal. Since this contribution is not spatially
homogeneous this contribution will lead to haze in the image, and,
therefore, reduce the dynamic range. It will also create image artefacts.
[0009] To avoid damage to the electronics and image artefacts due to scattered
radiation, an X-ray shield may be applied underneath the assembly of
scintillator or photoconductor and imaging array. Because of their high
density and high intrinsic stopping power for X-rays, metals with a high
atomic number are used as materials in such an X-ray shield. Examples of
these are sheets or plates from tantalum, lead or tungsten as disclosed in
EP1471384B1 , US2013/0032724A1 , US2012/0097857A1 .
[0010] However, metals with a high atomic number also have a high density.
Hence, X-ray shields based on these materials have a high weight. Weight
is an important characteristic of the RFPD especially for the portability of
the RFPDs. Any weight reduction is, therefore, beneficial for the users of
the RFPDs such as medical staff.
[001 1] US7317190B2 discloses a radiation absorbing X-ray detector panel
support comprising a radiation absorbing material to reduce the reflection
of X-rays of the back cover of the X-ray detector. The absorbing material
including heavy atoms such as lead, barium sulphate and tungsten can be
disposed as a film via a chemical vapour deposition technique onto a rigid
panel support or can be mixed via injection moulding with the base
materials used to fabricate the rigid panel support. The support for the
chemical vapour deposition as well as the base materials to fabricate the
rigid panel support, represent an extra weight contribution in the RFPD.
Moreover, the detector panel support comprising the radiation absorbing
material needs to be additionally fixed to assure immobilisation to the
detector.
[0012] In US5650626, an X-ray imaging detector is disclosed which contains a
substrate, supporting the conversion and detection unit. The substrate
includes one or more elements having atomic numbers greater than 22.
Since the detection array is directly deposited on the substrate, the variety
of suitable materials of the substrate is rather limited.
[0013] In US5777335, an imaging device is disclosed comprising a substrate,
preferably glass containing a metal selected from a group formed by Pb,
Ba, Ta or W. According to the inventors, the use of this glass would not
require an additional X-ray shield based on lead. However, glass
containing sufficient amounts of metals from a group formed by Pb, Ba, Ta
or is more expensive than glass which is normally used as a substrate
for imaging arrays.
[0014] US7569832 discloses a radiographic imaging device, namely a RFPD,
comprising two scintillating phosphor layers as scintillators each one
having different thicknesses and a transparent substrate to the X-rays
between said two layers. The use of an additional phosphor layer at the
opposite side of the substrate improves the X-ray absorption while
maintaining the spatial resolution. The presence of the additional phosphor
layer as disclosed is not sufficient to absorb all primary X-ray radiation to
prevent damage of the underlying electronics and to prevent backscatter.
An extra X-ray shield will still be required in the design of this RFPD.
[0015] In US2008/01 1960A1 a dual-screen digital radiography apparatus is
claimed. This apparatus consists of two flat panel detectors (front panel
and back panel) each comprising a scintillating phosphor layer to capture
and process X-rays. The scintillating phosphor layer in the back panel
contributes to the image formation and has no function as X-ray shield to
protect the underlying electronics. This dual-screen digital flat panel, still
requires an X-ray shield to protect the underlying electronics and to avoid
image artefacts due to scattered radiation.
[001 6] WO20051 055938 discloses a light weight film, with an X-ray absorption at
least equivalent to 0.254 mm of lead and which has to be applied on
garments or fabrics for personal radiation protection or attenuation, such
as aprons, thyroid shields, gonad shields, gloves, etc. Said film is obtained
from a polymer latex mixture comprising high atomic weight metals or their
related compounds and/or alloys. The suitable metals are the ones that
have an atomic number greater than 45. No use of this light weight film in
a RFPD is mentioned. Although a light weight film is claimed, the metal
particles used in the composition of the film still contribute to a high extend
to the weight of the shield.
[00 7] US6548570 discloses a radiation shielding composition to be applied on
garments or fabrics for personal radiation protection. The composition
comprised a polymer, preferably an elastomer, and a homogeneously
dispersed powder of a metal with high atomic number in an amount of at
least 80% in weight of the composition as filler. A loading material is mixed
with the filler material and kneaded with the elastomer at a temperature
below 180°C resulting in a radiation shielding composition that can be
applied homogeneously to garments and fabrics on an industrial scale.
The use of metals is however increasing the weight of the shield of this
invention considerably.
[0018] WO2009/0078891 discloses a radiation shielding sheet which is free from
lead and other harmful components having a highly radiation shielding
performance and an excellent economical efficiency. Said sheet is formed
by filling a shielding material into an organic polymer material, the
shielding material being an oxide powder containing at least one element
selected from the group consisting of lanthanum (La), cerium (Ce),
praseodymium (Pr), neodymium (Nd), samarium (Sm), europium (Eu) and
gadolinium (Gd) and the polymer being a material such as rubber,
thermoplastic elastomer, polymer resin or similar. The volumetric amount
of the shielding material filled in the radiation shielding sheet is 40 to 80
vol. % with respect to the total volume of the sheet. No use of this film into
a RFPD is mentioned.
[0019] From the foregoing discussion, it should be apparent that there is a need
for a RFPD with an X-ray shield to protect the underlying electronics and
to absorb the scattered radiation produced by the underlying structures to
avoid image artefacts in the imaging area, but which has a low weight, a
low cost, which can be produced in an economically efficient way and
which does not have to be fixed to the substrate of the imaging array in an
additional step of the production.
Summary of invention
[0020] It is therefore an object of the present invention to provide a solution for
the high weight contribution of the X-ray shield in a radiography flat panel
detector having a single imaging array and to provide at the same time a
solution for producing the RFPD on an economically efficient way. The
object has been achieved by a radiography flat panel detector as defined
in claim 1.
[0021] An additional advantage of the RFPD as defined in claiml , is that the
thickness of said X-ray shield can be adjusted in a continuous way to the
required degree of the X-ray shielding effect instead of in large steps as it
is in the case of shielding metal sheets commercially available with
standard thicknesses. Even though plates with custom made thickness
can be purchased, the price of those metal plates is still very high because
of the customization.
[0022] According to another aspect, the present invention includes a method of
manufacturing a radiography flat panel detector. The method includes
coating or depositing on the substrate of the imaging array, preferably on
the opposite side of the imaging array, an X-ray absorbing layer with at
least one chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements and which has
a dimensionless absorption exponent for 60 keV Am241 source greater
than 0.5 as defined in claim 1.
[0023] Other features, elements, steps, characteristics and advantages of the
present invention will become more apparent from the following detailed
description of preferred embodiments of the present invention. Specific
embodiments of the invention are also defined in the dependent claims.
Brief description of drawings
[0024] Fig. 1 represents a cross-section of a RFPD according to one embodiment
of the present invention and the underlying electronics, wherein:
1 is a scintillator or photoconductive layer
2 is single imaging array
3 is a substrate
4 is an X-ray absorbing layer
5 is the underlying electronics
Description of embodiments
[0025] The present invention relates to a radiography flat panel detector (RFPD)
comprising a scintillator or photoconductive layer, a single imaging array
on a substrate and an X-ray shield having an X-ray absorbing layer
comprising a chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements coated or
deposited on a side of a substrate of an imaging array. If the chemical
compound in the X-ray absorbing layer is a scintillating phosphor, a layer
is present between the X-ray absorbing layer and the substrate, which has
a transmission for light of 10% or lower at the wavelength of the light
emission of said chemical compound.
The X-ray absorbing layer
[0026] It has been found that X-ray shields can be made with the same X-ray
stopping power but with considerably less weight than X-ray shields
consisting of metals only by use of a layer comprising one or more
chemical compounds having a metal element with an atomic number of 20
or more and one or more non-metal elements. Preferably these
compounds are oxides or salts such as halides, oxysulphides, sulphites,
carbonates of metals with an atomic number of 20 or higher. Examples of
suitable metal elements with an atomic number higher than 20 that can be
used in the scope of the present invention are metals such as Barium (Ba),
Calcium (Ca), Cerium (Ce), Caesium (Cs), Gadolinium (Gd), Lanthanum
(La), Lutetium (Lu), Palladium (Pd), Tin (Sn), Strontium (Sr), Tellurium
(Te), Yttrium (Y), and Zinc (Zn). A further advantage of the invention is that
these compounds are relatively inexpensive and are characterised by a
low toxicity.
[0027] Examples of preferred compounds having a metal element with an atomic
number of 20 or more and one or more non-metal elements, are Caesium
iodide (Csl), Gadolinium oxysulphide (Gd202S), Barium fluorobromide
(BaFBr), Calcium tungstate (CaWCU), Barium titanate (BaTi0 3) ,
Gadolinium oxide (Gd203), Barium chloride (BaCb), Barium fluoride
(BaF2), Barium oxide (BaO), Cerium oxides, Caesium nitrate (CSNO3),
Gadolinium fluoride (GdF2) , Palladium iodide (Pdl2) , Tellurium dioxide
(Te0 2) , Tin iodides, Tin oxides, Barium sulphides, Barium carbonate
(BaC03), Barium iodide, Caesium chloride (CsCI), Caesium bromide
(CsBr), Caesium fluoride (CsF), Caesium sulphate (Cs2S0 4) , Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides, Rhenium sulphides, BaFX (wherein X represents CI or I), RFXn
(wherein RF represents lanthanides selected from: La, Ce, Pr, Nd, Pm,
Sm, Eu, Gd, Tb, Dy, Ho, Er, Tm, Yb, Lu and X represents halides selected
from: F, CI, Br, I), RFyOz, RFy(SO4)z, RFySz and/or RFy(WO4)z, wherein n,
y, z are independently an integer number higher than 1. These
compounds can produce lower weight X-ray shields and are easy to
handle due to their low hygroscopicity than their pure metal analogues.
The most preferred metallic compounds are: Gd2O2S, Gd2O3, Ce2O3, Csl,
BaFBr, CaWO and BaO.
[0028] It is another advantage of the present invention that the range of metal
elements which can be used for the x-ray absorbing layer, is much larger
than the corresponding range of the pure metals and/or alloys, since many
of them are not stable in their elemental form. Examples are the alkali
metals, the alkaline earth metals and the rare-earth metals.
[0029] The chemical compounds having a metal element with an atomic number
of 20 or more and one or more non-metal elements may be used in the Xray
absorbing layer of the present invention as powder dispersed in a
binder. The amount of the binder in the X-ray absorbing layer in weight
percent can vary in the range from 1% to 50%, preferably from 1% to 25%,
more preferably from 1% to 10%, most preferably from 1% to 3%.
[0030] Suitable binders are e.g. organic polymers or inorganic binding
components. Examples of suitable organic polymers are polyethylene
glycol acrylate, acrylic acid, butenoic acid, propenoic acid, urethane
acrylate, hexanediol diacrylate, copolyester tetracrylate, methylated
melamine, ethyl acetate, methyl methacrylate. Inorganic binding
components may be used as well. Examples of suitable inorganic binding
components are alumina, silica or alumina nanoparticles, aluminium
phosphate, sodium borate, barium phosphate, phosphoric acid, barium
nitrate.
[0031] Preferred binders are organic polymers such as cellulose acetate butyrate,
polyalkyl (meth)acrylates, polyvinyl-n-butyral, poly(vinylacetate-covinylchloride),
poly(acrylonitrile-co-butadiene-co-styrene), polyvinyl
chloride-co-vinyl acetate-co-vinylalcohol), poly(butyl acrylate), poly(ethyl
acrylate), poly(methacrylic acid), polyvinyl butyral), trimellitic acid,
butenedioic anhydride, phtalic anhydride, polyisoprene and/or a mixture
thereof. Preferably, the binder comprises one or more styrenehydrogenated
diene block copolymers, having a saturated rubber block
from polybutadiene or polyisoprene, as rubbery and/or elastomeric
polymers. Particularly suitable thermoplastic rubbers, which can be used
as block-copolymeric binders, in accordance with this invention, are the
KRATON™ G rubbers, KRATON™ being a trade name from SHELL.
[0032] In case the coating of the X-ray absorbing layer is to be cured, the binder
includes preferably a polymerisable compound which can be a
monofunctional or polyfunctional monomer, oligomer or polymer or a
combination thereof. The polymerisable compounds may comprise one or
more polymerisable groups, preferably radically polymerisable groups. Any
polymerisable mono- or oligofunctional monomer or oligomer commonly
known in the art may be employed. Preferred monofunctional monomers
are described in EP1637322A paragraph [0054] to [0057]. Preferred
oligofunctional monomers or oligomers are described in EP1637322A
paragraphs [0059] to [0064]. Particularly preferred polymerisable
compound are urethane (meth)acrylates and ,6-hexanedioldiacrylate.
The urethane (meth)acrylates are oligomer which may have one, two,
three or more polymerisable groups.
[0033] Suitable solvents, to dissolve the binder being an organic polymer during
the preparation of the coating solution of the X-ray absorbing layer can be
acetone, hexane, methyl acetate, ethyl acetate, isopropanol, methoxy
propanol, isobutyl acetate, ethanol, methanol, methylene chloride and
water. The most preferable ones are toluene, methyl-ethyl-ketone (MEK)
and methyl cyclohexane. To dissolve suitable inorganic binding
components, water is preferable as the main solvent. In case of a curable
coating liquid, one or more mono and/or difunctional monomers and/or
oligomers can be used as diluents. Preferred monomers and/or oligomers
acting as diluents are miscible with the above described urethane
(meth)acrylate oligomers. The monomer(s) or oligomer(s) used as diluents
are preferably low viscosity acrylate monomer(s).
[0034] The X-ray absorbing layer of the present invention may also comprise
additional compounds such as dispersants, plasticizers, photoinitiators,
photocurable monomers, antistatic agents, surfactants, stabilizers
oxidizing agents, adhesive agents, blocking agents and/or elastomers.
[0035] Dispersants which can be used in the present invention include nonsurface
active polymers or surface-active substances such as surfactants,
added to the binder to improve the separation of the particles of the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements and to further prevent
settling or clumping in the coating solution. Suitable examples of
dispersants are Stann JF95B from Sakyo and Disperse Ayd™ 1900 from
Daniel Produkts Company. The addition of dispersants to the coating
solution of the X-ray absorbing layer improves further the homogeneity of
the layer.
[0036] Suitable examples of plasticizers are Plastilit™ 3060 from BASF,
Santicizer™ 278 from Solutia Europe and Palatinol™ C from BASF. The
presence of plasticizers to the X-ray absorbing layer improves the
compatibility with flexible substrates.
[0037] Suitable photo-initiators are disclosed in e.g. J.V. Crivello et al. in "
Photoinitiators for Free Radical, Cationic & Anionic Photopolymerisation
2nd edition", Volume III of the Wiley/SITA Series In Surface Coatings
Technology, edited by G. Bradley and published in 1998 by John Wiley
and Sons Ltd London, pages 276 to 294.Examples of suitable
photoinitiators can be Darocure™ 1173 and Nuvopol™ PI-3000 from
Rahn. Examples of suitable antistatic agents can be Cyastat™ SN50 from
Acris and Lanco™ STAT K 100N from Langer.
[0038] Examples of suitable surfactants can be Dow Corning™ 190 and Gafac
RM710, Rhodafac™ RS-710 from Rodia. Examples of suitable stabilizer
compounds can be Brij™ 72 from ICI Surfactants and Barostab™ MS from
Baerlocher Italia. An example of a suitable oxidizing agent can be lead (IV)
oxide from Riedel De Haen. Examples of suitable adhesive agents can be
Craynor™ 435 from Cray Valley and Lanco™ wax TF1780 from Noveon.
An example of a suitable blocking agent can be Trixene™ BI7951 from
Baxenden. An example of a suitable elastomer compound can be Metaline
™from Schramm).
The thickness of the X-ray absorbing layer, the atomic number of the metal
element and the concentration of the chemical compound having a metal
element with an atomic number of 20 or more can be chosen to achieve a
desired level of X-ray absorption or attenuation in the RFPD. The value of
this level can be expressed as the "absorption exponent" (AE) and should
be equal to or higher than 0.5 to protect sufficiently the underlying
electronics of the RFPD and to limit the impact from backscattered X-rays
on the obtained image. The absorption exponent is a physical parameter
that is equal to the negative of the natural logarithm of the X-ray
transmittance. Since transmittance varies with X-ray energy, the
absorption exponent is more conveniently expressed relative to X-rays
emitted by a standard radiation source. A convenient standard is the 59.57
keV (hereafter 60 keV) gamma ray emission of Am241 . This source is in the
middle range of X-ray energies typically used in medical imaging, 20 to
150 keV, and is commonly used as a source of monoenergetic X-rays for
experiments. The absorption exponent can be measured directly or can be
calculated using formula 1 (expressed here for a 60 keV gamma ray
emission Am241 source):
AE( Am24 60keV)= t*(kiei +k2e2+k3e3+. ..) (Formula )
wherein AE(Am241 60 keV) represents the absorption exponent of the
substrate relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer in the principle
direction of propagation of the primary X-ray beam; e-i, b 2, 3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and
ki,k2,k3... represent the mass attenuation coefficients of the respective
elements at given energy. As the formula indicates, the absorption
exponent is equal to a thickness dimension multiplied by the sum of the
products of the mass attenuation coefficient for each element in the X-ray
absorbing layer at the about 60 keV gamma ray emission of Am241 and the
respective concentration of each element in the X-ray absorbing layer. The
absorption exponent is dimensionless. For example, if the mass
attenuation coefficients are expressed in cm2/mole, the concentrations
should be expressed in moles/cm3 and the thickness in centimetres. Mass
attenuation coefficients can be found on the 'National Institute for
Standards and Technology' (www.nist.gov/pml/data/xraycoef/). Depending
on the application, the coating weight of the chemical compound having a
metal element with an atomic number of 20 or more and one or more non
metal elements in the X-ray absorbing layer can be flexibly adjusted and in
case of using a RFPD for medical purposes, this coating weight is
preferably at least 100 mg/cm2, more preferably at least 200 mg/cm2.
[0040] The thickness of the X-ray absorbing layer can vary as well and depends
on the necessary shielding power and/or the space available to
incorporate the X-ray shield in the design of the RFPD. In the present
invention, the thickness of the X-ray absorbing layer can be at least 0.1
mm, more preferably in the range from 0.1 mm to 2.0 mm.
The light absorbing or light reflecting layer
[0041] Some of the chemical compounds having a metal element with an atomic
number of 20 or more and at least one non-metal elements are scintillating
phosphors which can emit light on X-ray absorption. If this is the case, light
emitted by these scintillating phosphors in the X-ray absorbing layer can
reach the imaging array through the substrate and contribute to the image
formation. Due to scattering in the substrate of the imaging array of the
light emitted by the scintillating phosphor present in the X-ray absorbing
layer, the quality of the image of the investigated object is negatively
impacted. In the case that scintillating phosphors are present in the X-ray
absorbing layer, a light reflecting or light absorbing layer is to be present
between the X-ray absorbing layer and the imaging array, more preferably
between the X-ray absorbing layer and the substrate of the imaging array.
In order to avoid any contribution of the emitted light by scintillating
phosphors in the X-ray absorbing layer to the image, the transmission of
the emitted light from the scintillating phosphor through this light absorbing
or reflecting layer, should be equal to or lower than 10%, more preferable
lower than 3%, most preferably lower than 1%. The term 'scintillating
phosphor' in the X-ray absorbing layer according to the invention should
be interpreted as a compound whose light emission on X-ray absorption
can reach the imaging array and contribute to the image formation of the
detector.
[0042] White coloured layers may be used to reflect light emitted by the
scintillating phosphor in the X-ray absorbing layer. Layers comprising T1O2
are preferably used to reflect 90% or more light at the wavelength(s) of the
light emitted by the scintillating phosphor. The solid content of T1O2 in the
light reflecting layer is preferably in the range of 25 to 50 (wt.)%. and the
thickness is preferably in the range of 5 to 40 pm. More preferably, the
solid content of the T1O2 is 33 to 38(wt.)% of the total solid content of the
layer and the layer thickness is between 13 and 30 pm. The layer is
preferably applied with a doctor blade coater on the substrate of the
imaging array, preferably on the side opposite to the imaging array.
[0043] In another preferred embodiment of the invention, black coloured layers
can be used to absorb light emitted by a scintillating phosphor in the X-ray
absorbing layer because of their high efficiency to absorb light. Black
particles, such as fine carbon black powder (ivory black, titanium black,
iron black), are suitable to obtain sufficient absorption of emitted light by
the scintillating phosphor. Preferably the solid content of carbon black is in
the range of 3 to 30 (wt.)% and a layer thickness of 2 to 30 pm will absorb
90% or more of the emitted light by the scintillating phosphor. More
preferably the range of the solid content of the carbon black is in the range
of 6 to 15 (wt.)% and the layer thickness between 5 and15 pm. In another
embodiment of the invention, coloured pigments or dyes absorbing
specifically at the maximal wavelength of the emitted light by the
scintillating phosphor in the X-ray absorbing layer can be used.
The scintillator
[0044] In the RFPD for indirect conversion direct radiography according to the
present invention, the scintillator comprises optionally a support and
provided thereon, a scintillating phosphor such as one or more of
Gd20 2S:Tb, Gd20 2S:Eu, Gd203:Eu, La20 2S:Tb, La20 2S, Y20 2S:Tb,
Cs TI, Csl:Eu, Csl:Na, CsBr:TI, Nal:TI, CaW0 , CaW0 :Tb, BaFBnEu,
BaFCLEu, BaS0 4:Eu, BaSrS0 4, BaPbS0 4, BaAh20i9 :Mn,
BaMgAlioOi 7:Eu, Zn2Si0 :Mn, (Zn, Cd)S:Ag, LaOBr, LaOBr:Tm,
Lu2O2S:Eu, Lu20 2S:Tb, LuTa04, Hf0 2:Ti, HfGe0 :Ti, YTa04, YTa04:Gd,
YTa04:Nb, Y203:Eu, YB0 3:Eu, YB0 3:Tb, or (Y,Gd)B0 3:Eu, or
combinations thereof. Besides crystalline scintillating phosphors,
scintillating glass or organic scintillators can also be used.
[0045] When evaporated under appropriate conditions, a layer of doped Csl will
condense in the form of needle like, closely packed crystallites with high
packing density onto a support. Such a columnar or needle-like scintillating
phosphor is known in the art. See, for example, ALN Stevels et al. , "Vapor
Deposited Csl:Na Layers: Screens for Application in X-Ray Imaging
Devices, " Philips Research Reports 29:353-362 (1974); and T. Jing et al,
"Enhanced Columnar Structure in Csl Layer by Substrate Patterning",
IEEE Trans. Nucl. Sci. 39: 1 95- 198 (1992). More preferably, the
scintillating phosphor layer includes doped Csl.
[0046] A blend of different scintillating phosphors can also be used. The median
particle size is generally between about 0. 5 m ti and about 40 mih . A
median particle size of between 1 mi and about 20 mih is preferred for
ease of formulation, as well as optimizing properties, such as speed,
sharpness and noise. The scintillator for the embodiments of the present
invention can be prepared using conventional coating techniques whereby
the scintillating phosphor powder, for example Gd20 2S is mixed with a
solution of a binder materia! and coated by means of a blade coater onto a
substrate. The binder can be chosen from a variety of known organic
polymers that are transparent to X-rays, stimulating, and emitting light.
Binders commonly employed in the art include sodium osulfobenzaldehyde
acetal of poly(vinyl alcohol); chloro-sulfonated
poly(ethylene); a mixture of macromolecular bisphenol poly(carbonates)
and copolymers comprising bisphenol carbonates and poly(alkylene
oxides);aqueous ethanol soluble nylons; poly(alkyl acrylates and
methacrylates) and copolymers of poly(alkyl acrylates and methacrylates
with acrylic and methacrylic acid); polyvinyl butyral); and poly(urethane)
elastomers. Other preferable binders which can be used are described
above in the section of the X-ray absorbing layer. Any conventional ratio
phosphor to binder can be employed. Generally, the thinner scintillating
phosphor layers are, the sharper images are realized when a high weight
ratio of phosphor to binder is employed. Phosphor-to-binder ratios in the
range of about 70:30 to 99:1 by weight are preferable.
The photoconductive layer
[0047] In the RFPD for direct conversion direct radiography according to the
present invention, the photoconductive layer is usually amorphous
selenium, although other photoconductors such as Hgl2, PbO, Pbl2, TIBr,
CdTe and gadolinium compounds can be used. The photoconductive layer
is preferentially deposited on the imaging array via vapour deposition but
can also been coated using any suitable coating method.
The imaging array and its substrate
[0048] The single imaging array used in the invention for indirect conversion
direct radiography is based on an indirect conversion process which uses
several physical components to convert X-rays into light that is
subsequently converted into electrical charges. First component is a
scintillating phosphor which converts X-rays into light (photons). Light is
further guided towards an amorphous silicon photodiode layer which
converts light into electrons and electrical charges are created. The
charges are collected and stored by the storage capacitors. A thin-film
transistor (TFT) array adjacent to amorphous silicon read out the electrical
charges and an image is created. Examples of suitable image arrays are
disclosed in US5262649 and by Samei E. et al., "General guidelines for
purchasing and acceptance testing of PACS equipment", Radiographics,
24, 313-334 . Preferably, the imaging arrays as described in
US201 3/0048866, paragraph [90-125] and US201 3/221 230, paragraphs
[53-71] and [81-104] can be used.
[0049] The imaging array used in the invention for direct conversion direct
radiography is based on a direct conversion process of X-ray photons into
electric charges. In this array, an electric field is created between a top
electrode, situated on top of the photoconductor layer and the TFT
elements. As X-rays strike the photoconductor, the electric charges are
created and the electrical field causes to move them towards the TFT
elements where they are collected and stored by storage capacitors.
Examples of suitable image arrays are disclosed by Samei E. et al., "
General guidelines for purchasing and acceptance testing of PACS
equipment", Radiographics, 24, 313-334.
[0050] For both the direct and indirect conversion process, the charges must be
read out by readout electronics. Examples of readout electronics in which
the electrical charges produced and stored are read out row by row, are
disclosed by Samei E. et al., Advances in Digital Radiography. RSNA
Categorical Course in Diagnostic Radiology Physics (p. 49-61) Oak Brook,
III.
[0051] The substrate of the imaging array of the present invention is preferably
glass. However, imaging arrays fabricated on substrates made of plastics,
metal foils can also be used. The imaging array can be protected from
humidity and environmental factors by a layer of silicon nitride or polymer
based coatings such as fluoropolymers, polyimides, polyamides,
polyurethanes and epoxy resins. Also polymers based on B-staged
bisbenzocyclobutene-based (BCB) monomers can be used. Alternatively,
porous inorganic dielectrics with low dielectric constants can also be used.
The underlying electronics
[0052] The underlying electronics, situated under the X-ray absorbing layer
comprise a circuit board which is equipped with electronic components for
processing the electrical signal from the imaging array, and/or controlling
the driver of the imaging array and is electrically connected to the imaging
array.
Method of making the radiographic flat panel detector
Method of making the X-ray shield
[0053] The X-ray shield of the present invention can be obtained by applying an
X-ray absorbing layer comprising at least one chemical compound having
a metal element with an atomic number of 20 or more and one or more
non-metal elements onto the substrate carrying the single imaging array.
Preferably, the X-ray absorbing layer is applied on the side of the
substrate opposite to the imaging array. Any known method for applying
layers on a substrate can be suitable, e.g. Physical Vapour Deposition
(PVD), Chemical Vapour Deposition (CVD), sputtering, doctor blade
coating, spin-coating, dip-coating, spray-coating, knife coating, screen
printing and lamination. The most preferable methods are doctor blade
coating and PVD.
[0054] One of the preferred methods of applying a layer is by coating a solution ,
hereafter denoted as coating solution, comprising the chemical compound
having a metal element with an atomic number of 20 or more and one or
more non-metal elements and a binder onto the substrate of the single
imaging array. In a preferred embodiment the coating solution is prepared
by first dissolving the binder in a suitable solvent. To this solution the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements is added. To obtain a
homogenous coating solution, a homogenization step or milling step of the
mixture can be included in the preparation process. A dispersant can be
added to the binder solution prior to the mixing with the chemical
compound having a metal element with an atomic number of 20 or more
and one or more non-metal elements. The dispersant improves the
separation of the particles in the coating solution and prevents settling or
clumping of the ingredients in the coating solution. The addition of
dispersants to the coating solution of the X-ray absorbing layer further
decreases the surface tension of the coating solution and improves the
coating quality of the X-ray absorbing layer.
[0055] In another embodiment of the invention, the binder being a polymerisable
compound can be dissolved in diluents comprising one or more mono
and/or difunctional monomers and/or oligomers.
[0056] After stirring or homogenization the coating solution is applied onto the
substrate preferably using a coating knife or a doctor blade. By adjusting
the distance between the coating blade and the substrate. After the
coating of the X-ray absorbing layer, this layer can be dried via an IRsource,
an UV-source, a heated metal roller or heated air. When
photocurable monomers are used in the coating solution, the coated layer
can be cured via heating or via an UV-source.
[0057] In another preferred embodiment, a PVD process is used in which the Xray
absorbing layer comprising the chemical compound having a metal
element with an atomic number of 20 or more and one or more non-metal
elements is prepared in vacuum from the gas phase of melting materials.
The material in a solid form can be introduced in a heat resistive container
to a vacuum chamber and subsequently heated to the temperature equal
to or higher than the melting point of compound(s). The melted material
vaporizes and condenses onto the substrate of the imaging array to form
the X-ray absorbing layer. Metal compounds as salts, halides, sulphides
and sulphates can be suitable in the PVD process due to their lower
melting point. The X-ray absorbing layer is than a deposited crystalline film
of chemical compound(s) having a metal element with an atomic number
of 20 or more and one or more non-metal elements and is binder-less.
[0058] It is an advantage of the present method of the invention that the X-ray
absorbing layer acting as an X-ray shield is directly applied on the
substrate of the imaging array. Hence, a step wherein the X-ray shield has
to be fixed to the substrate of the imaging array in the production, is
avoided.
[0059] In another embodiment, the X-ray absorbing layer can be applied on any
functional layer which was directly applied or coated on the substrate of
the imaging array prior to the application of the X-ray absorbing layer.
Examples of functional layers are: light absorbing layer, reflecting layer,
adhesion improving layer, protective layer, etc. Especially if the X-ray
absorbing layer comprises a scintillating phosphor, a layer is present
between the X-ray absorbing layer and the substrate, which has a
transmission for light of 10% or lower at the wavelength of the light
emission of the scintillating phosphor. This light absorbing or light
reflecting layer can be coated on the substrate of the imaging array using
conventional coating techniques known in the art.
Method of making the RFPD for indirect conversion direct radiography
[0060] The RFPD for indirect conversion direct radiography according to the
invention is made by assembling the different components which are
described above. A preferred method is now described.
[0061] After applying the X-ray absorbing layer on the substrate of the single
imaging array, the scintillator, which comprises a scintillating phosphor and
optionally a support, is coupled via gluing onto the imaging array. Gluing is
done with pressure sensitive adhesives or hot melts. Preferably a hot melt
is used. Suitable examples of hot melts are polyethylene-vinyl acetate,
polyolefins, polyamides, polyesters, polyurethanes, styrene block
copolymers, polycarbonates, fluoropolymers, silicone rubbers, polypyrrole.
The most preferred ones are polyolefins and polyurethanes due to the
higher temperature resistance and stability. The hot melt is preferably
thinner than 25 m. The hot melt with a lining is placed onto the surface of
the imaging array. The imaging array on its substrate, together with the hot
melt is then heated in an oven at a prescribed temperature. After cooling,
the lining is removed and releases a melted hot melt with a free adhesive
side. The scintillator is coupled to the imaging array by bringing the
scintillating phosphor layer in contact with the adhesive side of the hot melt
and by applying a high pressure at a high temperature. To achieve a good
sticking over the complete area of the imaging array, a pressure in a range
from 0.6 to 20 bars has to be applied and a temperature value in a range
from 80 - 220°C, during between 10 and 1000 s is required. A stack of
scintillator-imaging array-substrate- X-ray absorbing layer is thereby
formed.
[0062] In one preferred embodiment of the invention, this stack can be positioned
above the underlying electronics which perform the processing of the
electrical signal from the imaging array, or the controlling of the driver of
the imaging array.
[0063] In a preferred embodiment of the invention, the scintillator phosphor of the
scintillator is directly applied on the single imaging array via a coating or
deposition process. This method has the advantage that no gluing is
required and hence omits at least one step in the production process of
the RFPD. Another advantage of the direct application of the scintillating
phosphor on the imaging array, is the improved image quality.
[0064] In another embodiment of the invention, the X-ray absorbing layer is
applied to the substrate carrying the single imaging array, after the
scintillator has been coupled to the imaging array according to the
methods described above.
Method of making the RFPD for direct conversion direct radiography
[0065] The FPD for direct conversion direct radiography according to the
invention is made by assembling the different components which are
described above.
[0066] A preferred method is as follows: after applying the X-ray absorbing layer
to the substrate carrying the imaging array according to the same methods
as described for making the X-ray shield, the photoconductor, preferably
amorphous selenium is deposited onto the imaging array. Examples of
deposition methods are disclosed in Fischbach et al.,'Comparison of
indirect Csl/a:Si and direct a:Se digital radiography', Acta Radiologica 44
(2003) 616-621 . A top electrode on top of the photoconductive layer is
finally provided.
Examples
1. Method of measurement of the X-ray absorption:
1.1 . X-ray absorption measurement of the X-ray shields
[0067] The combination of the X-ray absorbing layer, the substrate and the
imaging array is denoted hereafter as X-ray shield. The X-ray absorption
of the X-ray shields was measured with a Philips Optimus 80 apparatus
together with Triad dosimeter having a 30cc volume cell. The X-ray shield
was placed with the imaging array directed towards the X-ray source. The
measuring cell was placed at 1.5 m distance from the X-ray source directly
behind the X-ray absorbing layer. All tests were done for standard
radiation X-ray beam qualities (RQA5 X-ray beam qualities as defined in
IEC standard 61267, 1st Ed. (1994)): RQA5 (21 mm Al, 73kV).
1.2. X-ray absorption measurement of the RFPD
[0068] RFPDs were produced by applying Gd202S or Csl scintillating phosphors
on the front side of the imaging array with its substrate having an X-ray
absorbing layer at the opposite side of the imaging array. The RFPD was
placed inside an in house-made frame, made of aluminium having a
thickness of 500 miti . The X-ray absorption of the RFPD was measured
with a Philips Optimus 80 apparatus together with Triad dosimeter having
a 30cc volume cell. The RFPD was placed with the scintillator directed
towards the X-ray source. The measuring cell was placed at 1.5 m
distance from the X-ray source directly behind the X-ray absorbing layer.
Data for each RFPD were collected multiple times and the average value
was calculated together with the standard deviation.
[0069] All tests were done for standard radiation X-ray beam qualities (RQA X-ray
beam qualities as defined in IEC standard 61267, 1st Ed. (1994)): RQA5
(21 mm Al, 73kV) and RQA9 (40 mm Al, 17kV).
2. Materials
[0070] Most materials used in the following examples were readily available from
standard sources such as ALDRICH CHEMICAL Co. (Belgium), ACROS
(Belgium) and BASF (Belgium) unless otherwise specified. All materials
were used without further purification unless otherwise specified.
• Gadolinium oxysulphide (Gd2O2S) or GOS: (CAS 12339-07-0) powder
was obtained from Nichia, mean particle size: 3.3 miti ;
• Caesium iodide (Csl): (CAS 7789-17-5) from Rockwood Lithium,
99.999%.
• Thl: Thallium iodide (CAS 62140-21-0) from Rockwood Lithium.
• Disperse Ayd™ 9 00 (Disperse Ayd™ W-22), anionic surfactant/Fatty
Ester dispersant (from Daniel Produkts Company).
• Kraton™ FG1901X (new name = Kraton™ FG1901 GT), a clear, linear
triblock copolymer based on styrene and ethylene/butylene with a
polystyrene content of 30%, from Shell Chemicals.
• Imaging array: TFT (according US201 3/0048866, paragraph [90-125]
and US201 3/221230, paragraphs [53-71] and [81-104]) on Corning
Lotus™ Glass substrate having a thickness of 0.7 mm and a size of
18cm X 24 cm.
• Aluminium having a thickness of 0.5 mm was obtained from Alanod.
• TiO2 R900:Ti-Pure ® R-900 Titanium Dioxide from DuPont.
• Filter AU09E1 1NG with pore size of 20 m from 3M.
• CAB 381-2: 20(wt.)% solution of Cellulose Acetate Butyrate (CAB-381-
2) from Eastman in MEK. Prepared by stirring for 8 hours at 1600 rpm
and filtering with Filter AU09E1 1NG after stirring.
Baysilone: Baysilone Paint additive MA from Bayer.
Ebecryl: 20(wt.)% solution of Ebecryl 1290, a hexafunctional aliphatic
urethane acrylate oligomer from Allnex in MEK, prepared by stirring for
8 hours at 1600 rpm and filtering with Filter AU09E1 1NG after stirring.
• Carbon black: Carbon black FW200 from Degussa
3. Preparation of X-ray shields
3.1 . Preparation of the solution for the coating of the X-ray absorbing layer
[0071] 4.5 g of binder (Kraton™ FG1901X) was dissolved in 18 g of a solvent
mixture of toluene and MEK (ratio 75:25 (wt.)) and stirred for 15 min at a
rate of 1900 r.p.m. The GOS was added thereafter in an amount of 200g
and the mixture was stirred for another 30 minutes at a rate of 1900 r.p.m.
The obtained GOS : binder ratio is 97.8 : 2.2 (wt).
3.2. Preparation of the solution for the light reflecting layer
[0072] 0.2 g of CAB 381-2 was mixed with 1 g of TiO2 R900, 0.001 g of Baysilone
and 2.6 g of MEK in a horizontal agitator bead mill. Finally Ebecryl was
added to achieve a CAB 381-2 : Ebecryl ratio of 1 : 1 (wt.). The solution
was filtered with Filter AU09E1 1NG. The solid content of TiO2 R900 is of
35(wt.)%.
3.3. Preparation of the solution for the light absorbing layer
[0073] 0.094 g of the 20 (wt.)% solution of CAB 381-2 in MEK as obtained in §
3.2., was mixed with 0.126 g of Carbon black, 0.001 g of Baysilone, 0.094
g of Ebecryl, and 3.686 g of MEK in a pearl mill (pearls: YTZ 0.8mm
diameter) for at least 30 min. The solid content of the Carbon black
obtained is 7.9 (wt.)%.
3.4. Preparation of X-ray shields SD-01 to SD-04 (INV) with GOS:
[0074] First the light reflecting layer was coated. The coating solution as obtained
in § 3.2. was coated with a doctor blade at a coating speed of 1.4 cm/s
onto the glass substrate of the imaging array on the side opposite to the
imaging array. The wet layer thickness was 250pm as to obtain a dry layer
thickness of 29 m. The drying of the light reflecting layer was done at
room temperature for at least 15 min. The transmission was measured at a
wavelength 550nm which correspond to the wavelength of the emitted
light by the scintillating phosphor GOS. The transmission value at 550nm
amounts to 5.2 %.
[0075] The coating solution as obtained in § 3.1 . was then coated with a doctor
blade at a coating speed of 4 m/min onto the previously coated light
reflecting layer. Different dry layer thicknesses variable from 100 to 450
m were obtained by adjusting the distance between the coating blade
and the substrate. Subsequently, the X-ray absorbing layer was dried at
room temperature during 30 minutes. In order to remove volatile solvents
as much as possible the coated X-ray shields were dried at 60°C for 30
minutes and again at 90°C for 20 to 30 minutes in a drying oven. The total
thickness of the X-ray absorbing layer was controlled by adjusting the wet
layer thickness and/or the number of layers coated on top of each other
after drying each layer. The wet layer thickness has a value between 220
pm and 1500 m.
[0076] After coating, each imaging array with the X-ray shield was weighed and
the coating weight of the chemical compound having a metal element with
an atomic number of 20 or more and one or more non metal elements in
the X-ray absorbing layer was obtained by applying formula 2. The results
are reported in Table 1
W, - W )
— *P%
Formula 2
Where:
F is the weight of the imaging array + substrate + X-ray absorbing layer,
Ws is the weight of the imaging array + substrate,
As is the surface area of the substrate,
P% is the amount in weight % of the chemical compound having a metal
element with an atomic number of 20 or more and one or more non-metal
elements in the X-ray absorbing layer.
3.5. Preparation of X-ray shield SD-05 (INV) with caesium iodide (Csl) :
SD-05 was prepared via physical vapour deposition of Csl on the
substrate of the imaging array. 400g of Csl was placed in a container in a
vacuum deposition chamber. The pressure in the chamber was decreased
to 5.10-5 mbar. The container was subsequently heated to a temperature
of 680°C and the Csl was deposited on the glass substrate on the side
opposite to the imaging array. The Csl-layer as obtained did not show a
substantial scintillating effect and hence can not be considered as a
phosphor scintillator. Indeed, only a very low light emission is observed
below 400 nm which is in a wavelength range where the imaging array is
not sensitive enough to contribute to the image of the investigated object.
The X-ray absorbing layer of Csl as obtained does not comprise a
scintillating phosphor and hence no light absorbing or light reflecting layer
was present between the substrate of the imaging array and the X-ray
absorbing layer comprising the Csl. The distance between the container
and the substrate was 20 cm. During evaporation, the substrate was
rotated at 12 r.p.m. and kept at elevated temperature of 140°C. During the
evaporation process argon gas was introduced into the chamber. The
duration of the process is 160 min. After the deposition, the imaging array
with its substrate and the X-ray shield was weighed and the coating weight
was obtained by applying formula 2 where P% is 100. The result is
reported in Table 1.
3.6. Molybdenum X-ray shield (COMP)
An X-ray shield based on a plate of Molybdenum (Mo) was obtained from
one of the commercially available RFPDs on the market. The thickness of
the Molybdenum plate was 0.3 mm. The Molybdenum plate did not contain
a substrate. The composition of the plate was 99.85% (wt.) of Mo, and
below 0.05% (wt.) of Na, K, Ca, Ni, Cu, and Bi.
The coating weight for this Mo plate was calculated based on formula 2
taking into account that F is the weight of the plate, P% is 00 and Ws is
0. The results of the calculated coating weight of the Mo plate, were
reported in Table 1.
Table 1: Coating weights and absorption exponent (AE) of the GOS or Csl
in the inventive X-ray shields (SD-01 to SD-05) and of the comparative Mo
plate.
Table 1
X-ray Compound having a metal Thickness Coating Absorption
shield element with an atomic of the X-ray weight exponent
number > 20 and > 1 non- absorbing (mg/cm2) (AE)
metal elements in the X-ray layer ( m)
absorbing layer
SD-01 GOS 325 172 0.79
(INV)
SD-02 GOS 325 72 0.79
(INV)
SD-03 GOS 230 115 0.56
(INV)
SD-04 GOS 330 155 0.80
(INV)
SD-05 Csl 300 112 0.56
(INV)
Mo-plate - 300 302 0.97
3.7. Preparation of X-ray shields with or without dispersant.
[0080] To illustrate the difference between GOS X-ray shields prepared with or
without a dispersant in the coating solution of the X-ray absorbing layer,
two X-ray shields based on GOS were prepared according to the method
described in §3.1 . Shield SD-01 was prepared without dispersant in the
coating solution and SD-02 was prepared with dispersant added to the
coating solution: 0.5 g of dispersant (Disperse Ayd™ 9100) was dissolved
in 11.21 g of a toluene and methyl-ethyl-ketone (MEK) solvent mixture,
having a ratio of 75:25 (wt) and mixed with the binder solution as prepared
in §3.1 . The further preparation steps are the same as described in §3.1 to
§3.4. The coating weight of the GOS was for both X-ray shields equal to
172mg/cm2. The X-ray absorption of both shields was determined
according § 1.1 . The results are shown in Table 2.
Table 2: X-ray absorption of GOS X-ray shields prepared with or without
dispersant.
Table 2
[0081] As shown in Table 2, the X-ray shield prepared with the dispersant present
in the coating solution had a more homogeneous X-ray absorbing layer for
a comparable weight and X-ray absorption as the X-ray shield prepared
without dispersant. The presence of the dispersant is advantageous for
the preparation process of the shields since it further reduces the surface
tension and prevents the floating of miti size particles.
4. X-ray absorption of inventive X-ray shields and comparative Mo shield
coupled to the substrate of the imaging array.
[0082] The X-ray absorption of the inventive X-ray shields SD-03, SD-05 and
comparative shield SD-06 was measured according § 1.1 . The comparative
X-ray shield SD-06 was obtained by contacting the Mo plate to the
substrate of the imaging array at the opposite side of the imaging array.
The results are shown in Table 3.
Table 3: Properties of inventive and comparative X-ray shields.
Table 3
[0083] Although the X-ray absorption of the inventive X-ray shields is lower than
the X-ray absorption of the comparative X-ray shield, the weight of the
inventive shields is considerably lower than the comparative X-ray shield.
Indeed, to have an absorption exponent for X-ray energies in the middle
range of X-ray energies typically used in medical imaging equal to X-ray
shield SD-05, the thickness of the Mo plate should be 170 m and hence
weigh considerably higher than SD-05. Unfortunately, Mo-plates with a
thickness of 70 m are not available and could hence not be included in
the example. The comparison of the two preferred compounds in the X-ray
absorbing layer of the inventive X-ray shields showed no significant
difference in the X-ray absorption capabilities.
5. Example 1
5.1 . Preparation of RFPDs comprising different X-ray shields
[0084] RFPDs for indirect conversion direct radiography were prepared by
bringing a scintillator in contact with the X-ray shields described in §3. To
assure a good optical contact between scintillating phosphor layer and the
imaging array, the scintillating phosphor was directly deposited or coated
on the imaging array. The scintillating phosphors used are GOS or needlebased
doped Csl. The GOS comprising scintillating phosphor layer was
prepared as follows: 0.5 g of dispersant (Disperse Ayd™ 9100) was
dissolved in 11.21 g of a toluene and methyl-ethyl-ketone (MEK) solvent
mixture, having a ratio of 75:25 (w/w) and mixed with the binder solution
as prepared in §3.1 . The obtained coating solution was coated on the
imaging array, the same way as §3.4. with a coating weight of 115
mg/cm 2. The needle-based doped Csl was prepared and deposited at a
coating weight of 120 mg/cm2 on the imaging array in the same way as
described in §3.5. with additional 1 (wt.)% of thallium dopant. The doping
with thallium was obtained by adding Thl to the Csl during the vapour
deposition process. The comparative RFPD, DRGOS-06 was prepared as
described above, but the X-ray absorbing layer on the substrate carrying
the imaging array is replaced by a Mo plate which was brought in contact
to the substrate of the imaging array at the opposite side of the imaging
array. The obtained RFPDs are summarised in Table 4.
Table 4: RFPDs based on different scintillators and X-ray shields.
Table 4
RFPD Scintillator X-ray shield
DRGOS-OI(INV) GOS SD-01
DRGOS-02(INV) GOS SD-02
DRGOS-03(INV) GOS SD-03
DRGOS-04(INV) GOS SD-04
DRGOS-05(INV) GOS SD-05
DRCS!-OI(INV) Csl SD-01
DRCSI-02(INV) Csl SD-02
DRCSI-03(INV) Csl SD-03
DRCSI-04(INV) Csl SD-04
DRCSI-05(INV) Csl SD-05
DRGOS-06 (COMP) GOS Mo
DRCSI-06 (COMP) Csl Mo
5.1. X-ray absorption of inventive and comparative RFPDs.
[0085] The X-ray absorption of inventive RFPDs (DRGOS-03 and DRGOS-04)
and a comparative RFPD (DRGOS-06) was measured according to § 1.2.
with following X-ray beam qualities and loads: RQA5 - 6.3 mAs and RQA9
- 3 mAs. The results of the measurements are provided in Table 5 .
Table 5: X-ray absorption of inventive and comparative RFPDs.
Table 5
[0086] The inventive RFPDs (DRGOS-03 and DRGOS-04) showed lower
absorption for X-ray beam quality RQA5 (6.3 mAs) in comparison with the
comparative RFPD (DRGOS-06). With the X-ray beam quality RQA9 (3
mAs), the inventive RFPDs (DRGOS-03 and DRGOS-04) showed a
comparable X-ray absorption as to the RFPD with the comparative Mo Xray
shield. The inventive RFPDs have, as additional advantage, a lower
weight than the comparative one. The inventive RFPDs can also be
produced on a more economically efficient way than the comparative one
since the fixing or gluing step between the substrate of the imaging array
and the X-ray absorbing layer is not required.
Claims
. A radiography flat panel detector comprising a layer configuration in the order
given,
a) a scintillator or photoconductive layer (1)
b) a single imaging array (2)
c) a substrate (3)
d) an X-ray absorbing layer (4) comprising a chemical compound having a
metal element with an atomic number of 20 or more and one or more nonmetal
elements,
characterised in that the X-ray absorbing layer has a dimensionless absorption
exponent of greater than 0.5 for gamma ray emission of Am241 at about 60keV;
wherein
AE( A m 24 60keV)= t* (kiei+k 2 e 2+k3 e3+ . . . )
wherein AE(Am241 60 keV) represents the absorption exponent of the X-ray
absorbing layer relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer; e-i , e2, e3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and ki,k 2 ,k3. . .
represent the mass attenuation coefficients of the respective elements, and if
the chemical compound is a scintillating phosphor, a layer is present between
the X-ray absorbing layer and the substrate, the layer having a transmission for
light of 10% or lower at the wavelength of the light emission of the chemical
compound.
2 . The radiography flat panel detector according to claim 1, wherein the X-ray
absorbing layer (4) is positioned between the substrate (3) and the underlying
electronics (5).
3. The radiography flat panel detector according to claim 1 or 2, wherein the
chemical compound is selected form the group consisting of Csl, Gd20 2S,
BaFBr, CaW0 , BaTiO3, Gd20 3, BaCI2, BaF2, BaO, Ce20 3, Ce0 2, CsN0 3
GdF2, Pdl2, Te0 2, Snl2, SnO, BaS0 , BaCOs, Bal, BaFX, RFXn, RFyO ,
RFy(S04)z, RFySz, RFy(W0 )z, CsBr, CsCI, CsF, CsN0 3, Cs2S0 Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides and Rhenium sulphides or mixtures thereof, wherein:
X is a halide selected from the group of F, CI, Br and I ; and
RF is a lanthanide selected from La, Ce, Pr, Nd, Pm, Sm, Eu, Gd,
Tb, Dy, Ho, Er, Tm, Yb and Lu; and
n, y, z are independently an integer number higher than .
4. The radiography flat panel detector according to any of the preceding claims,
wherein the X-ray absorbing layer comprises a binder.
5. The radiography flat panel detector according to claim 4, wherein the amount
of the binder in the X-ray absorbing layer is 10% by weight or less.
6 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, comprises a dye
or a pigment.
7. The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, is light absorbing.
8 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light emitted by the chemical
compound of 10% or lower, comprises light reflecting particles.
9 . A method of making a radiography flat panel detector as defined in claim 1,
comprising the steps of:
a) providing a substrate (3) with an imaging array (2) on a side of the first
substrate; and
b) applying a scintillating phosphor (1) onto the imaging array; and
c) applying the X-ray absorbing layer (4) on the side of the substrate opposite
to the imaging array.
10. The method of making a radiography flat panel detector according to claim 9
wherein the X-ray absorbing layer is coated by means of knife coating or
doctor blade coating.designed to convert the attenuated beam to a usable shadow image of the
internal structure of the object.
[0003] Increasingly, radiography flat panel detectors (RFPDs) are being used to
capture images of objects during inspection procedures or of body parts of
patients to be analyzed. These detectors can convert the X-rays directly
into electric charges (direct conversion direct radiography - DCDR), or in
an indirect way (indirect conversion direct radiography - ICDR).
[0004] In direct conversion direct radiography, the RFPDs convert X-rays directly
into electric charges. The X-rays are directly interacting with a
photoconductive layer such as amorphous selenium (a-Se).
[0005] In indirect conversion direct radiography, the RFPDs have a scintillating
phosphor such as Csl:TI (caesium iodide doped with thallium) or Gd20 2S
(gadolinium oxysulphide) which converts X-rays into light which then
interacts with an amorphous silicon (a-Si) semiconductor layer, where
electric charges are created.
[0006] The created electric charges are collected via a switching array,
comprising thin film transistors (TFTs). The transistors are switched-on
row by row and column by column to read out the signal of the detector.
The charges are transformed into voltage, which is converted in a digital
number that is stored in a computer file which can be used to generate a
softcopy or hardcopy image. Recently Complementary Metal Oxides
Semiconductors (CMOS) sensors are becoming important in X-ray
imaging. The detectors based on CMOS are already used in
mammography, dental, fluoroscopy, cardiology and angiography images.
The advantage of using those detectors is a high readout speed and a low
electronic noise.
[0007] Generally, the imaging array including TFTs as switching array and
photodiodes (in case of ICDR) is deposited on a thin substrate of glass.
The assembly of scintillator or photoconductor and the imaging array on
the glass substrate does not absorb all primary radiation, coming from the
X-ray source and transmitted by the object of the diagnosis. Hence the
electronics positioned under this assembly are exposed to a certain
fraction of the primary X-ray radiation. Since the electronics are not
sufficiently radiation hard, this transmitted radiation may cause damage.
[0008] Moreover, X-rays which are not absorbed by the assembly of scintillator or
photoconductor and the imaging array on the glass substrate, can be
absorbed in the structures underneath the glass substrate. The primary
radiation absorbed in these structures generates secondary radiation that
is emitted isotropically and that thus exposes the imaging part of the
detector. The secondary radiation is called "backscatter" and can expose
the image part of the detector thereby introducing artefacts into the
reconstructed image. Since the space under the assembly is not
homogeneously filled, the amount of scattered radiation is position
dependent. Part of the scattered radiation is emitted in the direction of the
assembly of scintillator or photoconductor and imaging array and may
contribute to the recorded signal. Since this contribution is not spatially
homogeneous this contribution will lead to haze in the image, and,
therefore, reduce the dynamic range. It will also create image artefacts.
[0009] To avoid damage to the electronics and image artefacts due to scattered
radiation, an X-ray shield may be applied underneath the assembly of
scintillator or photoconductor and imaging array. Because of their high
density and high intrinsic stopping power for X-rays, metals with a high
atomic number are used as materials in such an X-ray shield. Examples of
these are sheets or plates from tantalum, lead or tungsten as disclosed in
EP1471384B1 , US2013/0032724A1 , US2012/0097857A1 .
[0010] However, metals with a high atomic number also have a high density.
Hence, X-ray shields based on these materials have a high weight. Weight
is an important characteristic of the RFPD especially for the portability of
the RFPDs. Any weight reduction is, therefore, beneficial for the users of
the RFPDs such as medical staff.
[001 1] US7317190B2 discloses a radiation absorbing X-ray detector panel
support comprising a radiation absorbing material to reduce the reflection
of X-rays of the back cover of the X-ray detector. The absorbing material
including heavy atoms such as lead, barium sulphate and tungsten can be
disposed as a film via a chemical vapour deposition technique onto a rigid
panel support or can be mixed via injection moulding with the base
materials used to fabricate the rigid panel support. The support for the
chemical vapour deposition as well as the base materials to fabricate the
rigid panel support, represent an extra weight contribution in the RFPD.
Moreover, the detector panel support comprising the radiation absorbing
material needs to be additionally fixed to assure immobilisation to the
detector.
[0012] In US5650626, an X-ray imaging detector is disclosed which contains a
substrate, supporting the conversion and detection unit. The substrate
includes one or more elements having atomic numbers greater than 22.
Since the detection array is directly deposited on the substrate, the variety
of suitable materials of the substrate is rather limited.
[0013] In US5777335, an imaging device is disclosed comprising a substrate,
preferably glass containing a metal selected from a group formed by Pb,
Ba, Ta or W. According to the inventors, the use of this glass would not
require an additional X-ray shield based on lead. However, glass
containing sufficient amounts of metals from a group formed by Pb, Ba, Ta
or is more expensive than glass which is normally used as a substrate
for imaging arrays.
[0014] US7569832 discloses a radiographic imaging device, namely a RFPD,
comprising two scintillating phosphor layers as scintillators each one
having different thicknesses and a transparent substrate to the X-rays
between said two layers. The use of an additional phosphor layer at the
opposite side of the substrate improves the X-ray absorption while
maintaining the spatial resolution. The presence of the additional phosphor
layer as disclosed is not sufficient to absorb all primary X-ray radiation to
prevent damage of the underlying electronics and to prevent backscatter.
An extra X-ray shield will still be required in the design of this RFPD.
[0015] In US2008/01 1960A1 a dual-screen digital radiography apparatus is
claimed. This apparatus consists of two flat panel detectors (front panel
and back panel) each comprising a scintillating phosphor layer to capture
and process X-rays. The scintillating phosphor layer in the back panel
contributes to the image formation and has no function as X-ray shield to
protect the underlying electronics. This dual-screen digital flat panel, still
requires an X-ray shield to protect the underlying electronics and to avoid
image artefacts due to scattered radiation.
[001 6] WO20051 055938 discloses a light weight film, with an X-ray absorption at
least equivalent to 0.254 mm of lead and which has to be applied on
garments or fabrics for personal radiation protection or attenuation, such
as aprons, thyroid shields, gonad shields, gloves, etc. Said film is obtained
from a polymer latex mixture comprising high atomic weight metals or their
related compounds and/or alloys. The suitable metals are the ones that
have an atomic number greater than 45. No use of this light weight film in
a RFPD is mentioned. Although a light weight film is claimed, the metal
particles used in the composition of the film still contribute to a high extend
to the weight of the shield.
[00 7] US6548570 discloses a radiation shielding composition to be applied on
garments or fabrics for personal radiation protection. The composition
comprised a polymer, preferably an elastomer, and a homogeneously
dispersed powder of a metal with high atomic number in an amount of at
least 80% in weight of the composition as filler. A loading material is mixed
with the filler material and kneaded with the elastomer at a temperature
below 180°C resulting in a radiation shielding composition that can be
applied homogeneously to garments and fabrics on an industrial scale.
The use of metals is however increasing the weight of the shield of this
invention considerably.
[0018] WO2009/0078891 discloses a radiation shielding sheet which is free from
lead and other harmful components having a highly radiation shielding
performance and an excellent economical efficiency. Said sheet is formed
by filling a shielding material into an organic polymer material, the
shielding material being an oxide powder containing at least one element
selected from the group consisting of lanthanum (La), cerium (Ce),
praseodymium (Pr), neodymium (Nd), samarium (Sm), europium (Eu) and
gadolinium (Gd) and the polymer being a material such as rubber,
thermoplastic elastomer, polymer resin or similar. The volumetric amount
of the shielding material filled in the radiation shielding sheet is 40 to 80
vol. % with respect to the total volume of the sheet. No use of this film into
a RFPD is mentioned.
[0019] From the foregoing discussion, it should be apparent that there is a need
for a RFPD with an X-ray shield to protect the underlying electronics and
to absorb the scattered radiation produced by the underlying structures to
avoid image artefacts in the imaging area, but which has a low weight, a
low cost, which can be produced in an economically efficient way and
which does not have to be fixed to the substrate of the imaging array in an
additional step of the production.
Summary of invention
[0020] It is therefore an object of the present invention to provide a solution for
the high weight contribution of the X-ray shield in a radiography flat panel
detector having a single imaging array and to provide at the same time a
solution for producing the RFPD on an economically efficient way. The
object has been achieved by a radiography flat panel detector as defined
in claim 1.
[0021] An additional advantage of the RFPD as defined in claiml , is that the
thickness of said X-ray shield can be adjusted in a continuous way to the
required degree of the X-ray shielding effect instead of in large steps as it
is in the case of shielding metal sheets commercially available with
standard thicknesses. Even though plates with custom made thickness
can be purchased, the price of those metal plates is still very high because
of the customization.
[0022] According to another aspect, the present invention includes a method of
manufacturing a radiography flat panel detector. The method includes
coating or depositing on the substrate of the imaging array, preferably on
the opposite side of the imaging array, an X-ray absorbing layer with at
least one chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements and which has
a dimensionless absorption exponent for 60 keV Am241 source greater
than 0.5 as defined in claim 1.
[0023] Other features, elements, steps, characteristics and advantages of the
present invention will become more apparent from the following detailed
description of preferred embodiments of the present invention. Specific
embodiments of the invention are also defined in the dependent claims.
Brief description of drawings
[0024] Fig. 1 represents a cross-section of a RFPD according to one embodiment
of the present invention and the underlying electronics, wherein:
1 is a scintillator or photoconductive layer
2 is single imaging array
3 is a substrate
4 is an X-ray absorbing layer
5 is the underlying electronics
Description of embodiments
[0025] The present invention relates to a radiography flat panel detector (RFPD)
comprising a scintillator or photoconductive layer, a single imaging array
on a substrate and an X-ray shield having an X-ray absorbing layer
comprising a chemical compound having a metal element with an atomic
number of 20 or more and one or more non-metal elements coated or
deposited on a side of a substrate of an imaging array. If the chemical
compound in the X-ray absorbing layer is a scintillating phosphor, a layer
is present between the X-ray absorbing layer and the substrate, which has
a transmission for light of 10% or lower at the wavelength of the light
emission of said chemical compound.
The X-ray absorbing layer
[0026] It has been found that X-ray shields can be made with the same X-ray
stopping power but with considerably less weight than X-ray shields
consisting of metals only by use of a layer comprising one or more
chemical compounds having a metal element with an atomic number of 20
or more and one or more non-metal elements. Preferably these
compounds are oxides or salts such as halides, oxysulphides, sulphites,
carbonates of metals with an atomic number of 20 or higher. Examples of
suitable metal elements with an atomic number higher than 20 that can be
used in the scope of the present invention are metals such as Barium (Ba),
Calcium (Ca), Cerium (Ce), Caesium (Cs), Gadolinium (Gd), Lanthanum
(La), Lutetium (Lu), Palladium (Pd), Tin (Sn), Strontium (Sr), Tellurium
(Te), Yttrium (Y), and Zinc (Zn). A further advantage of the invention is that
these compounds are relatively inexpensive and are characterised by a
low toxicity.
[0027] Examples of preferred compounds having a metal element with an atomic
number of 20 or more and one or more non-metal elements, are Caesium
iodide (Csl), Gadolinium oxysulphide (Gd202S), Barium fluorobromide
(BaFBr), Calcium tungstate (CaWCU), Barium titanate (BaTi0 3) ,
Gadolinium oxide (Gd203), Barium chloride (BaCb), Barium fluoride
(BaF2), Barium oxide (BaO), Cerium oxides, Caesium nitrate (CSNO3),
Gadolinium fluoride (GdF2) , Palladium iodide (Pdl2) , Tellurium dioxide
(Te0 2) , Tin iodides, Tin oxides, Barium sulphides, Barium carbonate
(BaC03), Barium iodide, Caesium chloride (CsCI), Caesium bromide
(CsBr), Caesium fluoride (CsF), Caesium sulphate (Cs2S0 4) , Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides, Rhenium sulphides, BaFX (wherein X represents CI or I), RFXn
(wherein RF represents lanthanides selected from: La, Ce, Pr, Nd, Pm,
Sm, Eu, Gd, Tb, Dy, Ho, Er, Tm, Yb, Lu and X represents halides selected
from: F, CI, Br, I), RFyOz, RFy(SO4)z, RFySz and/or RFy(WO4)z, wherein n,
y, z are independently an integer number higher than 1. These
compounds can produce lower weight X-ray shields and are easy to
handle due to their low hygroscopicity than their pure metal analogues.
The most preferred metallic compounds are: Gd2O2S, Gd2O3, Ce2O3, Csl,
BaFBr, CaWO and BaO.
[0028] It is another advantage of the present invention that the range of metal
elements which can be used for the x-ray absorbing layer, is much larger
than the corresponding range of the pure metals and/or alloys, since many
of them are not stable in their elemental form. Examples are the alkali
metals, the alkaline earth metals and the rare-earth metals.
[0029] The chemical compounds having a metal element with an atomic number
of 20 or more and one or more non-metal elements may be used in the Xray
absorbing layer of the present invention as powder dispersed in a
binder. The amount of the binder in the X-ray absorbing layer in weight
percent can vary in the range from 1% to 50%, preferably from 1% to 25%,
more preferably from 1% to 10%, most preferably from 1% to 3%.
[0030] Suitable binders are e.g. organic polymers or inorganic binding
components. Examples of suitable organic polymers are polyethylene
glycol acrylate, acrylic acid, butenoic acid, propenoic acid, urethane
acrylate, hexanediol diacrylate, copolyester tetracrylate, methylated
melamine, ethyl acetate, methyl methacrylate. Inorganic binding
components may be used as well. Examples of suitable inorganic binding
components are alumina, silica or alumina nanoparticles, aluminium
phosphate, sodium borate, barium phosphate, phosphoric acid, barium
nitrate.
[0031] Preferred binders are organic polymers such as cellulose acetate butyrate,
polyalkyl (meth)acrylates, polyvinyl-n-butyral, poly(vinylacetate-covinylchloride),
poly(acrylonitrile-co-butadiene-co-styrene), polyvinyl
chloride-co-vinyl acetate-co-vinylalcohol), poly(butyl acrylate), poly(ethyl
acrylate), poly(methacrylic acid), polyvinyl butyral), trimellitic acid,
butenedioic anhydride, phtalic anhydride, polyisoprene and/or a mixture
thereof. Preferably, the binder comprises one or more styrenehydrogenated
diene block copolymers, having a saturated rubber block
from polybutadiene or polyisoprene, as rubbery and/or elastomeric
polymers. Particularly suitable thermoplastic rubbers, which can be used
as block-copolymeric binders, in accordance with this invention, are the
KRATON™ G rubbers, KRATON™ being a trade name from SHELL.
[0032] In case the coating of the X-ray absorbing layer is to be cured, the binder
includes preferably a polymerisable compound which can be a
monofunctional or polyfunctional monomer, oligomer or polymer or a
combination thereof. The polymerisable compounds may comprise one or
more polymerisable groups, preferably radically polymerisable groups. Any
polymerisable mono- or oligofunctional monomer or oligomer commonly
known in the art may be employed. Preferred monofunctional monomers
are described in EP1637322A paragraph [0054] to [0057]. Preferred
oligofunctional monomers or oligomers are described in EP1637322A
paragraphs [0059] to [0064]. Particularly preferred polymerisable
compound are urethane (meth)acrylates and ,6-hexanedioldiacrylate.
The urethane (meth)acrylates are oligomer which may have one, two,
three or more polymerisable groups.
[0033] Suitable solvents, to dissolve the binder being an organic polymer during
the preparation of the coating solution of the X-ray absorbing layer can be
acetone, hexane, methyl acetate, ethyl acetate, isopropanol, methoxy
propanol, isobutyl acetate, ethanol, methanol, methylene chloride and
water. The most preferable ones are toluene, methyl-ethyl-ketone (MEK)
and methyl cyclohexane. To dissolve suitable inorganic binding
components, water is preferable as the main solvent. In case of a curable
coating liquid, one or more mono and/or difunctional monomers and/or
oligomers can be used as diluents. Preferred monomers and/or oligomers
acting as diluents are miscible with the above described urethane
(meth)acrylate oligomers. The monomer(s) or oligomer(s) used as diluents
are preferably low viscosity acrylate monomer(s).
[0034] The X-ray absorbing layer of the present invention may also comprise
additional compounds such as dispersants, plasticizers, photoinitiators,
photocurable monomers, antistatic agents, surfactants, stabilizers
oxidizing agents, adhesive agents, blocking agents and/or elastomers.
[0035] Dispersants which can be used in the present invention include nonsurface
active polymers or surface-active substances such as surfactants,
added to the binder to improve the separation of the particles of the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements and to further prevent
settling or clumping in the coating solution. Suitable examples of
dispersants are Stann JF95B from Sakyo and Disperse Ayd™ 1900 from
Daniel Produkts Company. The addition of dispersants to the coating
solution of the X-ray absorbing layer improves further the homogeneity of
the layer.
[0036] Suitable examples of plasticizers are Plastilit™ 3060 from BASF,
Santicizer™ 278 from Solutia Europe and Palatinol™ C from BASF. The
presence of plasticizers to the X-ray absorbing layer improves the
compatibility with flexible substrates.
[0037] Suitable photo-initiators are disclosed in e.g. J.V. Crivello et al. in "
Photoinitiators for Free Radical, Cationic & Anionic Photopolymerisation
2nd edition", Volume III of the Wiley/SITA Series In Surface Coatings
Technology, edited by G. Bradley and published in 1998 by John Wiley
and Sons Ltd London, pages 276 to 294.Examples of suitable
photoinitiators can be Darocure™ 1173 and Nuvopol™ PI-3000 from
Rahn. Examples of suitable antistatic agents can be Cyastat™ SN50 from
Acris and Lanco™ STAT K 100N from Langer.
[0038] Examples of suitable surfactants can be Dow Corning™ 190 and Gafac
RM710, Rhodafac™ RS-710 from Rodia. Examples of suitable stabilizer
compounds can be Brij™ 72 from ICI Surfactants and Barostab™ MS from
Baerlocher Italia. An example of a suitable oxidizing agent can be lead (IV)
oxide from Riedel De Haen. Examples of suitable adhesive agents can be
Craynor™ 435 from Cray Valley and Lanco™ wax TF1780 from Noveon.
An example of a suitable blocking agent can be Trixene™ BI7951 from
Baxenden. An example of a suitable elastomer compound can be Metaline
™from Schramm).
The thickness of the X-ray absorbing layer, the atomic number of the metal
element and the concentration of the chemical compound having a metal
element with an atomic number of 20 or more can be chosen to achieve a
desired level of X-ray absorption or attenuation in the RFPD. The value of
this level can be expressed as the "absorption exponent" (AE) and should
be equal to or higher than 0.5 to protect sufficiently the underlying
electronics of the RFPD and to limit the impact from backscattered X-rays
on the obtained image. The absorption exponent is a physical parameter
that is equal to the negative of the natural logarithm of the X-ray
transmittance. Since transmittance varies with X-ray energy, the
absorption exponent is more conveniently expressed relative to X-rays
emitted by a standard radiation source. A convenient standard is the 59.57
keV (hereafter 60 keV) gamma ray emission of Am241 . This source is in the
middle range of X-ray energies typically used in medical imaging, 20 to
150 keV, and is commonly used as a source of monoenergetic X-rays for
experiments. The absorption exponent can be measured directly or can be
calculated using formula 1 (expressed here for a 60 keV gamma ray
emission Am241 source):
AE( Am24 60keV)= t*(kiei +k2e2+k3e3+. ..) (Formula )
wherein AE(Am241 60 keV) represents the absorption exponent of the
substrate relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer in the principle
direction of propagation of the primary X-ray beam; e-i, b 2, 3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and
ki,k2,k3... represent the mass attenuation coefficients of the respective
elements at given energy. As the formula indicates, the absorption
exponent is equal to a thickness dimension multiplied by the sum of the
products of the mass attenuation coefficient for each element in the X-ray
absorbing layer at the about 60 keV gamma ray emission of Am241 and the
respective concentration of each element in the X-ray absorbing layer. The
absorption exponent is dimensionless. For example, if the mass
attenuation coefficients are expressed in cm2/mole, the concentrations
should be expressed in moles/cm3 and the thickness in centimetres. Mass
attenuation coefficients can be found on the 'National Institute for
Standards and Technology' (www.nist.gov/pml/data/xraycoef/). Depending
on the application, the coating weight of the chemical compound having a
metal element with an atomic number of 20 or more and one or more non
metal elements in the X-ray absorbing layer can be flexibly adjusted and in
case of using a RFPD for medical purposes, this coating weight is
preferably at least 100 mg/cm2, more preferably at least 200 mg/cm2.
[0040] The thickness of the X-ray absorbing layer can vary as well and depends
on the necessary shielding power and/or the space available to
incorporate the X-ray shield in the design of the RFPD. In the present
invention, the thickness of the X-ray absorbing layer can be at least 0.1
mm, more preferably in the range from 0.1 mm to 2.0 mm.
The light absorbing or light reflecting layer
[0041] Some of the chemical compounds having a metal element with an atomic
number of 20 or more and at least one non-metal elements are scintillating
phosphors which can emit light on X-ray absorption. If this is the case, light
emitted by these scintillating phosphors in the X-ray absorbing layer can
reach the imaging array through the substrate and contribute to the image
formation. Due to scattering in the substrate of the imaging array of the
light emitted by the scintillating phosphor present in the X-ray absorbing
layer, the quality of the image of the investigated object is negatively
impacted. In the case that scintillating phosphors are present in the X-ray
absorbing layer, a light reflecting or light absorbing layer is to be present
between the X-ray absorbing layer and the imaging array, more preferably
between the X-ray absorbing layer and the substrate of the imaging array.
In order to avoid any contribution of the emitted light by scintillating
phosphors in the X-ray absorbing layer to the image, the transmission of
the emitted light from the scintillating phosphor through this light absorbing
or reflecting layer, should be equal to or lower than 10%, more preferable
lower than 3%, most preferably lower than 1%. The term 'scintillating
phosphor' in the X-ray absorbing layer according to the invention should
be interpreted as a compound whose light emission on X-ray absorption
can reach the imaging array and contribute to the image formation of the
detector.
[0042] White coloured layers may be used to reflect light emitted by the
scintillating phosphor in the X-ray absorbing layer. Layers comprising T1O2
are preferably used to reflect 90% or more light at the wavelength(s) of the
light emitted by the scintillating phosphor. The solid content of T1O2 in the
light reflecting layer is preferably in the range of 25 to 50 (wt.)%. and the
thickness is preferably in the range of 5 to 40 pm. More preferably, the
solid content of the T1O2 is 33 to 38(wt.)% of the total solid content of the
layer and the layer thickness is between 13 and 30 pm. The layer is
preferably applied with a doctor blade coater on the substrate of the
imaging array, preferably on the side opposite to the imaging array.
[0043] In another preferred embodiment of the invention, black coloured layers
can be used to absorb light emitted by a scintillating phosphor in the X-ray
absorbing layer because of their high efficiency to absorb light. Black
particles, such as fine carbon black powder (ivory black, titanium black,
iron black), are suitable to obtain sufficient absorption of emitted light by
the scintillating phosphor. Preferably the solid content of carbon black is in
the range of 3 to 30 (wt.)% and a layer thickness of 2 to 30 pm will absorb
90% or more of the emitted light by the scintillating phosphor. More
preferably the range of the solid content of the carbon black is in the range
of 6 to 15 (wt.)% and the layer thickness between 5 and15 pm. In another
embodiment of the invention, coloured pigments or dyes absorbing
specifically at the maximal wavelength of the emitted light by the
scintillating phosphor in the X-ray absorbing layer can be used.
The scintillator
[0044] In the RFPD for indirect conversion direct radiography according to the
present invention, the scintillator comprises optionally a support and
provided thereon, a scintillating phosphor such as one or more of
Gd20 2S:Tb, Gd20 2S:Eu, Gd203:Eu, La20 2S:Tb, La20 2S, Y20 2S:Tb,
Cs TI, Csl:Eu, Csl:Na, CsBr:TI, Nal:TI, CaW0 , CaW0 :Tb, BaFBnEu,
BaFCLEu, BaS0 4:Eu, BaSrS0 4, BaPbS0 4, BaAh20i9 :Mn,
BaMgAlioOi 7:Eu, Zn2Si0 :Mn, (Zn, Cd)S:Ag, LaOBr, LaOBr:Tm,
Lu2O2S:Eu, Lu20 2S:Tb, LuTa04, Hf0 2:Ti, HfGe0 :Ti, YTa04, YTa04:Gd,
YTa04:Nb, Y203:Eu, YB0 3:Eu, YB0 3:Tb, or (Y,Gd)B0 3:Eu, or
combinations thereof. Besides crystalline scintillating phosphors,
scintillating glass or organic scintillators can also be used.
[0045] When evaporated under appropriate conditions, a layer of doped Csl will
condense in the form of needle like, closely packed crystallites with high
packing density onto a support. Such a columnar or needle-like scintillating
phosphor is known in the art. See, for example, ALN Stevels et al. , "Vapor
Deposited Csl:Na Layers: Screens for Application in X-Ray Imaging
Devices, " Philips Research Reports 29:353-362 (1974); and T. Jing et al,
"Enhanced Columnar Structure in Csl Layer by Substrate Patterning",
IEEE Trans. Nucl. Sci. 39: 1 95- 198 (1992). More preferably, the
scintillating phosphor layer includes doped Csl.
[0046] A blend of different scintillating phosphors can also be used. The median
particle size is generally between about 0. 5 m ti and about 40 mih . A
median particle size of between 1 mi and about 20 mih is preferred for
ease of formulation, as well as optimizing properties, such as speed,
sharpness and noise. The scintillator for the embodiments of the present
invention can be prepared using conventional coating techniques whereby
the scintillating phosphor powder, for example Gd20 2S is mixed with a
solution of a binder materia! and coated by means of a blade coater onto a
substrate. The binder can be chosen from a variety of known organic
polymers that are transparent to X-rays, stimulating, and emitting light.
Binders commonly employed in the art include sodium osulfobenzaldehyde
acetal of poly(vinyl alcohol); chloro-sulfonated
poly(ethylene); a mixture of macromolecular bisphenol poly(carbonates)
and copolymers comprising bisphenol carbonates and poly(alkylene
oxides);aqueous ethanol soluble nylons; poly(alkyl acrylates and
methacrylates) and copolymers of poly(alkyl acrylates and methacrylates
with acrylic and methacrylic acid); polyvinyl butyral); and poly(urethane)
elastomers. Other preferable binders which can be used are described
above in the section of the X-ray absorbing layer. Any conventional ratio
phosphor to binder can be employed. Generally, the thinner scintillating
phosphor layers are, the sharper images are realized when a high weight
ratio of phosphor to binder is employed. Phosphor-to-binder ratios in the
range of about 70:30 to 99:1 by weight are preferable.
The photoconductive layer
[0047] In the RFPD for direct conversion direct radiography according to the
present invention, the photoconductive layer is usually amorphous
selenium, although other photoconductors such as Hgl2, PbO, Pbl2, TIBr,
CdTe and gadolinium compounds can be used. The photoconductive layer
is preferentially deposited on the imaging array via vapour deposition but
can also been coated using any suitable coating method.
The imaging array and its substrate
[0048] The single imaging array used in the invention for indirect conversion
direct radiography is based on an indirect conversion process which uses
several physical components to convert X-rays into light that is
subsequently converted into electrical charges. First component is a
scintillating phosphor which converts X-rays into light (photons). Light is
further guided towards an amorphous silicon photodiode layer which
converts light into electrons and electrical charges are created. The
charges are collected and stored by the storage capacitors. A thin-film
transistor (TFT) array adjacent to amorphous silicon read out the electrical
charges and an image is created. Examples of suitable image arrays are
disclosed in US5262649 and by Samei E. et al., "General guidelines for
purchasing and acceptance testing of PACS equipment", Radiographics,
24, 313-334 . Preferably, the imaging arrays as described in
US201 3/0048866, paragraph [90-125] and US201 3/221 230, paragraphs
[53-71] and [81-104] can be used.
[0049] The imaging array used in the invention for direct conversion direct
radiography is based on a direct conversion process of X-ray photons into
electric charges. In this array, an electric field is created between a top
electrode, situated on top of the photoconductor layer and the TFT
elements. As X-rays strike the photoconductor, the electric charges are
created and the electrical field causes to move them towards the TFT
elements where they are collected and stored by storage capacitors.
Examples of suitable image arrays are disclosed by Samei E. et al., "
General guidelines for purchasing and acceptance testing of PACS
equipment", Radiographics, 24, 313-334.
[0050] For both the direct and indirect conversion process, the charges must be
read out by readout electronics. Examples of readout electronics in which
the electrical charges produced and stored are read out row by row, are
disclosed by Samei E. et al., Advances in Digital Radiography. RSNA
Categorical Course in Diagnostic Radiology Physics (p. 49-61) Oak Brook,
III.
[0051] The substrate of the imaging array of the present invention is preferably
glass. However, imaging arrays fabricated on substrates made of plastics,
metal foils can also be used. The imaging array can be protected from
humidity and environmental factors by a layer of silicon nitride or polymer
based coatings such as fluoropolymers, polyimides, polyamides,
polyurethanes and epoxy resins. Also polymers based on B-staged
bisbenzocyclobutene-based (BCB) monomers can be used. Alternatively,
porous inorganic dielectrics with low dielectric constants can also be used.
The underlying electronics
[0052] The underlying electronics, situated under the X-ray absorbing layer
comprise a circuit board which is equipped with electronic components for
processing the electrical signal from the imaging array, and/or controlling
the driver of the imaging array and is electrically connected to the imaging
array.
Method of making the radiographic flat panel detector
Method of making the X-ray shield
[0053] The X-ray shield of the present invention can be obtained by applying an
X-ray absorbing layer comprising at least one chemical compound having
a metal element with an atomic number of 20 or more and one or more
non-metal elements onto the substrate carrying the single imaging array.
Preferably, the X-ray absorbing layer is applied on the side of the
substrate opposite to the imaging array. Any known method for applying
layers on a substrate can be suitable, e.g. Physical Vapour Deposition
(PVD), Chemical Vapour Deposition (CVD), sputtering, doctor blade
coating, spin-coating, dip-coating, spray-coating, knife coating, screen
printing and lamination. The most preferable methods are doctor blade
coating and PVD.
[0054] One of the preferred methods of applying a layer is by coating a solution ,
hereafter denoted as coating solution, comprising the chemical compound
having a metal element with an atomic number of 20 or more and one or
more non-metal elements and a binder onto the substrate of the single
imaging array. In a preferred embodiment the coating solution is prepared
by first dissolving the binder in a suitable solvent. To this solution the
chemical compound having a metal element with an atomic number of 20
or more and one or more non-metal elements is added. To obtain a
homogenous coating solution, a homogenization step or milling step of the
mixture can be included in the preparation process. A dispersant can be
added to the binder solution prior to the mixing with the chemical
compound having a metal element with an atomic number of 20 or more
and one or more non-metal elements. The dispersant improves the
separation of the particles in the coating solution and prevents settling or
clumping of the ingredients in the coating solution. The addition of
dispersants to the coating solution of the X-ray absorbing layer further
decreases the surface tension of the coating solution and improves the
coating quality of the X-ray absorbing layer.
[0055] In another embodiment of the invention, the binder being a polymerisable
compound can be dissolved in diluents comprising one or more mono
and/or difunctional monomers and/or oligomers.
[0056] After stirring or homogenization the coating solution is applied onto the
substrate preferably using a coating knife or a doctor blade. By adjusting
the distance between the coating blade and the substrate. After the
coating of the X-ray absorbing layer, this layer can be dried via an IRsource,
an UV-source, a heated metal roller or heated air. When
photocurable monomers are used in the coating solution, the coated layer
can be cured via heating or via an UV-source.
[0057] In another preferred embodiment, a PVD process is used in which the Xray
absorbing layer comprising the chemical compound having a metal
element with an atomic number of 20 or more and one or more non-metal
elements is prepared in vacuum from the gas phase of melting materials.
The material in a solid form can be introduced in a heat resistive container
to a vacuum chamber and subsequently heated to the temperature equal
to or higher than the melting point of compound(s). The melted material
vaporizes and condenses onto the substrate of the imaging array to form
the X-ray absorbing layer. Metal compounds as salts, halides, sulphides
and sulphates can be suitable in the PVD process due to their lower
melting point. The X-ray absorbing layer is than a deposited crystalline film
of chemical compound(s) having a metal element with an atomic number
of 20 or more and one or more non-metal elements and is binder-less.
[0058] It is an advantage of the present method of the invention that the X-ray
absorbing layer acting as an X-ray shield is directly applied on the
substrate of the imaging array. Hence, a step wherein the X-ray shield has
to be fixed to the substrate of the imaging array in the production, is
avoided.
[0059] In another embodiment, the X-ray absorbing layer can be applied on any
functional layer which was directly applied or coated on the substrate of
the imaging array prior to the application of the X-ray absorbing layer.
Examples of functional layers are: light absorbing layer, reflecting layer,
adhesion improving layer, protective layer, etc. Especially if the X-ray
absorbing layer comprises a scintillating phosphor, a layer is present
between the X-ray absorbing layer and the substrate, which has a
transmission for light of 10% or lower at the wavelength of the light
emission of the scintillating phosphor. This light absorbing or light
reflecting layer can be coated on the substrate of the imaging array using
conventional coating techniques known in the art.
Method of making the RFPD for indirect conversion direct radiography
[0060] The RFPD for indirect conversion direct radiography according to the
invention is made by assembling the different components which are
described above. A preferred method is now described.
[0061] After applying the X-ray absorbing layer on the substrate of the single
imaging array, the scintillator, which comprises a scintillating phosphor and
optionally a support, is coupled via gluing onto the imaging array. Gluing is
done with pressure sensitive adhesives or hot melts. Preferably a hot melt
is used. Suitable examples of hot melts are polyethylene-vinyl acetate,
polyolefins, polyamides, polyesters, polyurethanes, styrene block
copolymers, polycarbonates, fluoropolymers, silicone rubbers, polypyrrole.
The most preferred ones are polyolefins and polyurethanes due to the
higher temperature resistance and stability. The hot melt is preferably
thinner than 25 m. The hot melt with a lining is placed onto the surface of
the imaging array. The imaging array on its substrate, together with the hot
melt is then heated in an oven at a prescribed temperature. After cooling,
the lining is removed and releases a melted hot melt with a free adhesive
side. The scintillator is coupled to the imaging array by bringing the
scintillating phosphor layer in contact with the adhesive side of the hot melt
and by applying a high pressure at a high temperature. To achieve a good
sticking over the complete area of the imaging array, a pressure in a range
from 0.6 to 20 bars has to be applied and a temperature value in a range
from 80 - 220°C, during between 10 and 1000 s is required. A stack of
scintillator-imaging array-substrate- X-ray absorbing layer is thereby
formed.
[0062] In one preferred embodiment of the invention, this stack can be positioned
above the underlying electronics which perform the processing of the
electrical signal from the imaging array, or the controlling of the driver of
the imaging array.
[0063] In a preferred embodiment of the invention, the scintillator phosphor of the
scintillator is directly applied on the single imaging array via a coating or
deposition process. This method has the advantage that no gluing is
required and hence omits at least one step in the production process of
the RFPD. Another advantage of the direct application of the scintillating
phosphor on the imaging array, is the improved image quality.
[0064] In another embodiment of the invention, the X-ray absorbing layer is
applied to the substrate carrying the single imaging array, after the
scintillator has been coupled to the imaging array according to the
methods described above.
Method of making the RFPD for direct conversion direct radiography
[0065] The FPD for direct conversion direct radiography according to the
invention is made by assembling the different components which are
described above.
[0066] A preferred method is as follows: after applying the X-ray absorbing layer
to the substrate carrying the imaging array according to the same methods
as described for making the X-ray shield, the photoconductor, preferably
amorphous selenium is deposited onto the imaging array. Examples of
deposition methods are disclosed in Fischbach et al.,'Comparison of
indirect Csl/a:Si and direct a:Se digital radiography', Acta Radiologica 44
(2003) 616-621 . A top electrode on top of the photoconductive layer is
finally provided.
Examples
1. Method of measurement of the X-ray absorption:
1.1 . X-ray absorption measurement of the X-ray shields
[0067] The combination of the X-ray absorbing layer, the substrate and the
imaging array is denoted hereafter as X-ray shield. The X-ray absorption
of the X-ray shields was measured with a Philips Optimus 80 apparatus
together with Triad dosimeter having a 30cc volume cell. The X-ray shield
was placed with the imaging array directed towards the X-ray source. The
measuring cell was placed at 1.5 m distance from the X-ray source directly
behind the X-ray absorbing layer. All tests were done for standard
radiation X-ray beam qualities (RQA5 X-ray beam qualities as defined in
IEC standard 61267, 1st Ed. (1994)): RQA5 (21 mm Al, 73kV).
1.2. X-ray absorption measurement of the RFPD
[0068] RFPDs were produced by applying Gd202S or Csl scintillating phosphors
on the front side of the imaging array with its substrate having an X-ray
absorbing layer at the opposite side of the imaging array. The RFPD was
placed inside an in house-made frame, made of aluminium having a
thickness of 500 miti . The X-ray absorption of the RFPD was measured
with a Philips Optimus 80 apparatus together with Triad dosimeter having
a 30cc volume cell. The RFPD was placed with the scintillator directed
towards the X-ray source. The measuring cell was placed at 1.5 m
distance from the X-ray source directly behind the X-ray absorbing layer.
Data for each RFPD were collected multiple times and the average value
was calculated together with the standard deviation.
[0069] All tests were done for standard radiation X-ray beam qualities (RQA X-ray
beam qualities as defined in IEC standard 61267, 1st Ed. (1994)): RQA5
(21 mm Al, 73kV) and RQA9 (40 mm Al, 17kV).
2. Materials
[0070] Most materials used in the following examples were readily available from
standard sources such as ALDRICH CHEMICAL Co. (Belgium), ACROS
(Belgium) and BASF (Belgium) unless otherwise specified. All materials
were used without further purification unless otherwise specified.
• Gadolinium oxysulphide (Gd2O2S) or GOS: (CAS 12339-07-0) powder
was obtained from Nichia, mean particle size: 3.3 miti ;
• Caesium iodide (Csl): (CAS 7789-17-5) from Rockwood Lithium,
99.999%.
• Thl: Thallium iodide (CAS 62140-21-0) from Rockwood Lithium.
• Disperse Ayd™ 9 00 (Disperse Ayd™ W-22), anionic surfactant/Fatty
Ester dispersant (from Daniel Produkts Company).
• Kraton™ FG1901X (new name = Kraton™ FG1901 GT), a clear, linear
triblock copolymer based on styrene and ethylene/butylene with a
polystyrene content of 30%, from Shell Chemicals.
• Imaging array: TFT (according US201 3/0048866, paragraph [90-125]
and US201 3/221230, paragraphs [53-71] and [81-104]) on Corning
Lotus™ Glass substrate having a thickness of 0.7 mm and a size of
18cm X 24 cm.
• Aluminium having a thickness of 0.5 mm was obtained from Alanod.
• TiO2 R900:Ti-Pure ® R-900 Titanium Dioxide from DuPont.
• Filter AU09E1 1NG with pore size of 20 m from 3M.
• CAB 381-2: 20(wt.)% solution of Cellulose Acetate Butyrate (CAB-381-
2) from Eastman in MEK. Prepared by stirring for 8 hours at 1600 rpm
and filtering with Filter AU09E1 1NG after stirring.
Baysilone: Baysilone Paint additive MA from Bayer.
Ebecryl: 20(wt.)% solution of Ebecryl 1290, a hexafunctional aliphatic
urethane acrylate oligomer from Allnex in MEK, prepared by stirring for
8 hours at 1600 rpm and filtering with Filter AU09E1 1NG after stirring.
• Carbon black: Carbon black FW200 from Degussa
3. Preparation of X-ray shields
3.1 . Preparation of the solution for the coating of the X-ray absorbing layer
[0071] 4.5 g of binder (Kraton™ FG1901X) was dissolved in 18 g of a solvent
mixture of toluene and MEK (ratio 75:25 (wt.)) and stirred for 15 min at a
rate of 1900 r.p.m. The GOS was added thereafter in an amount of 200g
and the mixture was stirred for another 30 minutes at a rate of 1900 r.p.m.
The obtained GOS : binder ratio is 97.8 : 2.2 (wt).
3.2. Preparation of the solution for the light reflecting layer
[0072] 0.2 g of CAB 381-2 was mixed with 1 g of TiO2 R900, 0.001 g of Baysilone
and 2.6 g of MEK in a horizontal agitator bead mill. Finally Ebecryl was
added to achieve a CAB 381-2 : Ebecryl ratio of 1 : 1 (wt.). The solution
was filtered with Filter AU09E1 1NG. The solid content of TiO2 R900 is of
35(wt.)%.
3.3. Preparation of the solution for the light absorbing layer
[0073] 0.094 g of the 20 (wt.)% solution of CAB 381-2 in MEK as obtained in §
3.2., was mixed with 0.126 g of Carbon black, 0.001 g of Baysilone, 0.094
g of Ebecryl, and 3.686 g of MEK in a pearl mill (pearls: YTZ 0.8mm
diameter) for at least 30 min. The solid content of the Carbon black
obtained is 7.9 (wt.)%.
3.4. Preparation of X-ray shields SD-01 to SD-04 (INV) with GOS:
[0074] First the light reflecting layer was coated. The coating solution as obtained
in § 3.2. was coated with a doctor blade at a coating speed of 1.4 cm/s
onto the glass substrate of the imaging array on the side opposite to the
imaging array. The wet layer thickness was 250pm as to obtain a dry layer
thickness of 29 m. The drying of the light reflecting layer was done at
room temperature for at least 15 min. The transmission was measured at a
wavelength 550nm which correspond to the wavelength of the emitted
light by the scintillating phosphor GOS. The transmission value at 550nm
amounts to 5.2 %.
[0075] The coating solution as obtained in § 3.1 . was then coated with a doctor
blade at a coating speed of 4 m/min onto the previously coated light
reflecting layer. Different dry layer thicknesses variable from 100 to 450
m were obtained by adjusting the distance between the coating blade
and the substrate. Subsequently, the X-ray absorbing layer was dried at
room temperature during 30 minutes. In order to remove volatile solvents
as much as possible the coated X-ray shields were dried at 60°C for 30
minutes and again at 90°C for 20 to 30 minutes in a drying oven. The total
thickness of the X-ray absorbing layer was controlled by adjusting the wet
layer thickness and/or the number of layers coated on top of each other
after drying each layer. The wet layer thickness has a value between 220
pm and 1500 m.
[0076] After coating, each imaging array with the X-ray shield was weighed and
the coating weight of the chemical compound having a metal element with
an atomic number of 20 or more and one or more non metal elements in
the X-ray absorbing layer was obtained by applying formula 2. The results
are reported in Table 1
W, - W )
— *P%
Formula 2
Where:
F is the weight of the imaging array + substrate + X-ray absorbing layer,
Ws is the weight of the imaging array + substrate,
As is the surface area of the substrate,
P% is the amount in weight % of the chemical compound having a metal
element with an atomic number of 20 or more and one or more non-metal
elements in the X-ray absorbing layer.
3.5. Preparation of X-ray shield SD-05 (INV) with caesium iodide (Csl) :
SD-05 was prepared via physical vapour deposition of Csl on the
substrate of the imaging array. 400g of Csl was placed in a container in a
vacuum deposition chamber. The pressure in the chamber was decreased
to 5.10-5 mbar. The container was subsequently heated to a temperature
of 680°C and the Csl was deposited on the glass substrate on the side
opposite to the imaging array. The Csl-layer as obtained did not show a
substantial scintillating effect and hence can not be considered as a
phosphor scintillator. Indeed, only a very low light emission is observed
below 400 nm which is in a wavelength range where the imaging array is
not sensitive enough to contribute to the image of the investigated object.
The X-ray absorbing layer of Csl as obtained does not comprise a
scintillating phosphor and hence no light absorbing or light reflecting layer
was present between the substrate of the imaging array and the X-ray
absorbing layer comprising the Csl. The distance between the container
and the substrate was 20 cm. During evaporation, the substrate was
rotated at 12 r.p.m. and kept at elevated temperature of 140°C. During the
evaporation process argon gas was introduced into the chamber. The
duration of the process is 160 min. After the deposition, the imaging array
with its substrate and the X-ray shield was weighed and the coating weight
was obtained by applying formula 2 where P% is 100. The result is
reported in Table 1.
3.6. Molybdenum X-ray shield (COMP)
An X-ray shield based on a plate of Molybdenum (Mo) was obtained from
one of the commercially available RFPDs on the market. The thickness of
the Molybdenum plate was 0.3 mm. The Molybdenum plate did not contain
a substrate. The composition of the plate was 99.85% (wt.) of Mo, and
below 0.05% (wt.) of Na, K, Ca, Ni, Cu, and Bi.
The coating weight for this Mo plate was calculated based on formula 2
taking into account that F is the weight of the plate, P% is 00 and Ws is
0. The results of the calculated coating weight of the Mo plate, were
reported in Table 1.
Table 1: Coating weights and absorption exponent (AE) of the GOS or Csl
in the inventive X-ray shields (SD-01 to SD-05) and of the comparative Mo
plate.
Table 1
X-ray Compound having a metal Thickness Coating Absorption
shield element with an atomic of the X-ray weight exponent
number > 20 and > 1 non- absorbing (mg/cm2) (AE)
metal elements in the X-ray layer ( m)
absorbing layer
SD-01 GOS 325 172 0.79
(INV)
SD-02 GOS 325 72 0.79
(INV)
SD-03 GOS 230 115 0.56
(INV)
SD-04 GOS 330 155 0.80
(INV)
SD-05 Csl 300 112 0.56
(INV)
Mo-plate - 300 302 0.97
3.7. Preparation of X-ray shields with or without dispersant.
[0080] To illustrate the difference between GOS X-ray shields prepared with or
without a dispersant in the coating solution of the X-ray absorbing layer,
two X-ray shields based on GOS were prepared according to the method
described in §3.1 . Shield SD-01 was prepared without dispersant in the
coating solution and SD-02 was prepared with dispersant added to the
coating solution: 0.5 g of dispersant (Disperse Ayd™ 9100) was dissolved
in 11.21 g of a toluene and methyl-ethyl-ketone (MEK) solvent mixture,
having a ratio of 75:25 (wt) and mixed with the binder solution as prepared
in §3.1 . The further preparation steps are the same as described in §3.1 to
§3.4. The coating weight of the GOS was for both X-ray shields equal to
172mg/cm2. The X-ray absorption of both shields was determined
according § 1.1 . The results are shown in Table 2.
Table 2: X-ray absorption of GOS X-ray shields prepared with or without
dispersant.
Table 2
[0081] As shown in Table 2, the X-ray shield prepared with the dispersant present
in the coating solution had a more homogeneous X-ray absorbing layer for
a comparable weight and X-ray absorption as the X-ray shield prepared
without dispersant. The presence of the dispersant is advantageous for
the preparation process of the shields since it further reduces the surface
tension and prevents the floating of miti size particles.
4. X-ray absorption of inventive X-ray shields and comparative Mo shield
coupled to the substrate of the imaging array.
[0082] The X-ray absorption of the inventive X-ray shields SD-03, SD-05 and
comparative shield SD-06 was measured according § 1.1 . The comparative
X-ray shield SD-06 was obtained by contacting the Mo plate to the
substrate of the imaging array at the opposite side of the imaging array.
The results are shown in Table 3.
Table 3: Properties of inventive and comparative X-ray shields.
Table 3
[0083] Although the X-ray absorption of the inventive X-ray shields is lower than
the X-ray absorption of the comparative X-ray shield, the weight of the
inventive shields is considerably lower than the comparative X-ray shield.
Indeed, to have an absorption exponent for X-ray energies in the middle
range of X-ray energies typically used in medical imaging equal to X-ray
shield SD-05, the thickness of the Mo plate should be 170 m and hence
weigh considerably higher than SD-05. Unfortunately, Mo-plates with a
thickness of 70 m are not available and could hence not be included in
the example. The comparison of the two preferred compounds in the X-ray
absorbing layer of the inventive X-ray shields showed no significant
difference in the X-ray absorption capabilities.
5. Example 1
5.1 . Preparation of RFPDs comprising different X-ray shields
[0084] RFPDs for indirect conversion direct radiography were prepared by
bringing a scintillator in contact with the X-ray shields described in §3. To
assure a good optical contact between scintillating phosphor layer and the
imaging array, the scintillating phosphor was directly deposited or coated
on the imaging array. The scintillating phosphors used are GOS or needlebased
doped Csl. The GOS comprising scintillating phosphor layer was
prepared as follows: 0.5 g of dispersant (Disperse Ayd™ 9100) was
dissolved in 11.21 g of a toluene and methyl-ethyl-ketone (MEK) solvent
mixture, having a ratio of 75:25 (w/w) and mixed with the binder solution
as prepared in §3.1 . The obtained coating solution was coated on the
imaging array, the same way as §3.4. with a coating weight of 115
mg/cm 2. The needle-based doped Csl was prepared and deposited at a
coating weight of 120 mg/cm2 on the imaging array in the same way as
described in §3.5. with additional 1 (wt.)% of thallium dopant. The doping
with thallium was obtained by adding Thl to the Csl during the vapour
deposition process. The comparative RFPD, DRGOS-06 was prepared as
described above, but the X-ray absorbing layer on the substrate carrying
the imaging array is replaced by a Mo plate which was brought in contact
to the substrate of the imaging array at the opposite side of the imaging
array. The obtained RFPDs are summarised in Table 4.
Table 4: RFPDs based on different scintillators and X-ray shields.
Table 4
RFPD Scintillator X-ray shield
DRGOS-OI(INV) GOS SD-01
DRGOS-02(INV) GOS SD-02
DRGOS-03(INV) GOS SD-03
DRGOS-04(INV) GOS SD-04
DRGOS-05(INV) GOS SD-05
DRCS!-OI(INV) Csl SD-01
DRCSI-02(INV) Csl SD-02
DRCSI-03(INV) Csl SD-03
DRCSI-04(INV) Csl SD-04
DRCSI-05(INV) Csl SD-05
DRGOS-06 (COMP) GOS Mo
DRCSI-06 (COMP) Csl Mo
5.1. X-ray absorption of inventive and comparative RFPDs.
[0085] The X-ray absorption of inventive RFPDs (DRGOS-03 and DRGOS-04)
and a comparative RFPD (DRGOS-06) was measured according to § 1.2.
with following X-ray beam qualities and loads: RQA5 - 6.3 mAs and RQA9
- 3 mAs. The results of the measurements are provided in Table 5 .
Table 5: X-ray absorption of inventive and comparative RFPDs.
Table 5
[0086] The inventive RFPDs (DRGOS-03 and DRGOS-04) showed lower
absorption for X-ray beam quality RQA5 (6.3 mAs) in comparison with the
comparative RFPD (DRGOS-06). With the X-ray beam quality RQA9 (3
mAs), the inventive RFPDs (DRGOS-03 and DRGOS-04) showed a
comparable X-ray absorption as to the RFPD with the comparative Mo Xray
shield. The inventive RFPDs have, as additional advantage, a lower
weight than the comparative one. The inventive RFPDs can also be
produced on a more economically efficient way than the comparative one
since the fixing or gluing step between the substrate of the imaging array
and the X-ray absorbing layer is not required.
Claims
. A radiography flat panel detector comprising a layer configuration in the order
given,
a) a scintillator or photoconductive layer (1)
b) a single imaging array (2)
c) a substrate (3)
d) an X-ray absorbing layer (4) comprising a chemical compound having a
metal element with an atomic number of 20 or more and one or more nonmetal
elements,
characterised in that the X-ray absorbing layer has a dimensionless absorption
exponent of greater than 0.5 for gamma ray emission of Am241 at about 60keV;
wherein
AE( A m 24 60keV)= t* (kiei+k 2 e 2+k3 e3+ . . . )
wherein AE(Am241 60 keV) represents the absorption exponent of the X-ray
absorbing layer relative to the about 60 keV gamma ray emission of Am241 ; t
represents the thickness of the X-ray absorbing layer; e-i , e2, e3, ... represent
the concentrations of the elements in the X-ray absorbing layer; and ki,k 2 ,k3. . .
represent the mass attenuation coefficients of the respective elements, and if
the chemical compound is a scintillating phosphor, a layer is present between
the X-ray absorbing layer and the substrate, the layer having a transmission for
light of 10% or lower at the wavelength of the light emission of the chemical
compound.
2 . The radiography flat panel detector according to claim 1, wherein the X-ray
absorbing layer (4) is positioned between the substrate (3) and the underlying
electronics (5).
3. The radiography flat panel detector according to claim 1 or 2, wherein the
chemical compound is selected form the group consisting of Csl, Gd20 2S,
BaFBr, CaW0 , BaTiO3, Gd20 3, BaCI2, BaF2, BaO, Ce20 3, Ce0 2, CsN0 3
GdF2, Pdl2, Te0 2, Snl2, SnO, BaS0 , BaCOs, Bal, BaFX, RFXn, RFyO ,
RFy(S04)z, RFySz, RFy(W0 )z, CsBr, CsCI, CsF, CsN0 3, Cs2S0 Osmium
halides, Osmium oxides, Osmium sulphides, Rhenium halides, Rhenium
oxides and Rhenium sulphides or mixtures thereof, wherein:
X is a halide selected from the group of F, CI, Br and I ; and
RF is a lanthanide selected from La, Ce, Pr, Nd, Pm, Sm, Eu, Gd,
Tb, Dy, Ho, Er, Tm, Yb and Lu; and
n, y, z are independently an integer number higher than .
4. The radiography flat panel detector according to any of the preceding claims,
wherein the X-ray absorbing layer comprises a binder.
5. The radiography flat panel detector according to claim 4, wherein the amount
of the binder in the X-ray absorbing layer is 10% by weight or less.
6 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, comprises a dye
or a pigment.
7. The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light of 10% or lower at the
wavelength of the light emission of the chemical compound, is light absorbing.
8 . The radiography flat panel detector according to any of the preceding claims,
wherein the layer having a transmission for the light emitted by the chemical
compound of 10% or lower, comprises light reflecting particles.
9 . A method of making a radiography flat panel detector as defined in claim 1,
comprising the steps of:
a) providing a substrate (3) with an imaging array (2) on a side of the first
substrate; and
b) applying a scintillating phosphor (1) onto the imaging array; and
c) applying the X-ray absorbing layer (4) on the side of the substrate opposite
to the imaging array.
10. The method of making a radiography flat panel detector according to claim 9
wherein the X-ray absorbing layer is coated by means of knife coating or
doctor blade coating.