Energy resolved X-ray diffraction Cl. J.Kosanetzky, G.Harding, U.Neitzel

Transcription

1 Proc. Of SPIE Vol 0626, Application of Optical Instrumentation in Medicine XIV and Picture Archiving and Communication Systems(PACS IV) for Medical Applications IV, ed. S J Dwyer/R H Schneider (Jan 1986) Energy resolved X-ray diffraction Cl J.Kosanetzky, G.Harding, U.Neitzel Philips GmbH Forschungslaboratorium Hamburg Vogt-Kdlln-Str. 30, 2000 Hamburg 54, W. Germany Abstract Low angle x-ray scattering at diagnostic energies in narrow beam geometry is due to coherent (Rayleigh) and incoherent (Compton) scattering. It has been found that single coherent scatter dominates below IO deg. Interference effects with coherent scatter leads to diffraction patterns which differ from material to material. A technique, analogous to conventional CT, allows the reconstruction of the 20 distributipn20f the x-ray diffraction properties within an object slice, as demonstrated recently 9. Use of the bremsstrahlung spectrum of an x-ray tube permits short measuring times, but causes a significant energy broadening of the diffraction curves, thus deteriorating the maximum contrast obtainable by diffraction imaging. With energy resolved photon counting of the scattered x-ray quanta this broadening can be corrected, yielding an image contrast approaching that of a monochromatic x-ray source. Introduction ---- Conventional x-ray CT techniques rely on the detection of transmitted x-ray quanta. The total attenuation of the incident radiation by the object is due to photoelectric absorption, Compton and Rayleigh scattering in the normal diagnostic energy range ( KeV). Elastic (Rayleigh) scatter contributes only 1X to the total attenuation and is therefore normally neglected. But several authors have demonstrated rtckently that elastic scatter dominates the low angle x-ray scatter below about IO deg. 3. The pattern resulting from interference phenomena of this elastic scattering are characteristic for the material radiated. This offers the chance for better tissue characterisation than with normal transmission CT. The angular positions of the measured diffraction maxima depend on the x-ray energy. Therefore, use of a bremsstrahlung spectrum from an x-ray tube results in short measuring times due to the high x-ray intensities but introduces an energy broadening of the scatter curves. We present a method to correct the measurements for this broadening by energy resolved photon counting. Method The aim of this method is to image a 20 section within an object by x-radiation. The measured parameter is the angular differential cross-section for elastic scatter da/dq. As in normal CT imaging the spatial distribution of this cross-section can be reconstructed from a complete set of measured line integrals (projections). Because da/dq is a function of the scatter angle, a number of different images can be reconstructed, representing the angular variation of elastic scatter for each pixel. Fig. 1 explains the basic concept of our imaging method. The radiation of an x-ray tube is collimated into a pencil beam which penetrates the object. Instead of only one detector for transmission measurements a number of detector elements are arranged on either side of the central detector, or slternatively a single detector is shifted step by step to measure the scattered x-ray intensity as a function of the scatter angle. The signal of the i-th detector can be written as: Si Z IO I Tp * n0 * (do/dq) * (dq/da) * TB dl with: IO = incident intensity TP = attenuation before scatter event 0 = number of molecules per unit volume dq/da q solid angle corresponding to the detector area TB = attenuation of scattered radiation. SPIE Vo/. 626 Medicine XIV/PACS IV/l 986J/ 137

2 To correct the measurements for attenuation along the beampath within the object we made the following assumption. Due to the small scatter angle the deviation of the scattered x-ray quanta from the primary beam path is rather small and we approximate the total attenuation of the scattered intensity by: Ty = Tp * TS with TT = attenuation of the transmitted beam. Therefore, normalisation against the transmitted intensity corrects for attenuation effects. Because we have a background of single Compton and multiple scatteri g we correct 3 our detector readings for this background according to Monte Carlo calculations. Reconstruction is done by a standard filtered backprojection algorithm, yielding a set of images at different scatter angles. Energy broadenineof the scatter curves --- The technique described so far yields the exact scatter curves only for monochromatic radiation. In the case of polyenergetic radiation the diffraction patterns at different energies are not correctly superposed. A pattern at energy EO will contract in angular space as the energy is increased to El according to: x = EO * sin(b0/2.) = El * sin(b1/2.) (3) with x = prop. to the momentum transfer 0 q scatter angle. Hence the use of a polyenergetic x-ray source leads to a blurring of the diffraction structure. The fine structure is washed out and it is now much more difficult to differentiate between materials, see Fig. 2. Use of a monochromatic source like a radio nuclide would remove this blurring but increase the measurement time considerably. Correction of the energy blurring Even with a polyenergetic source one can obtain sharp diffraction pattern as if measured with monochromatic radiation by the following procedure: Consider a specific detector position Di in Fig.1 which corresponds for a specified object element (voxel) to a certain scatter angle Bi. The energy of the incident x-ray quanta varies over the whole range of the tube, so that the diffraction intensity at position Di is the integral over a broad range of x values (eq. 3). The total diffraction pattern is a superposition of the patterns for each energy, see Fig. 3. If the detector is energy sensitive, it is possible to count the scattered quanta into different channels according to their individual energies El,EZ,E3 etc. Each energy channel corresponds to a specific x channel depending on the scatter angle which corresponds to the detector position. So one can reorder the counted x-ray quanta into x channels. After completion of the whole measurement a set of x projections is available for reconstruction providing the wanted variation of elastic scatter cross-section for each pixel. Experimentalstem and results ---- We speculated earlier that the diffraction properties of materials may allow better characterisation and detection than use of transmitted radiation alone. To validate this speculation, we performed a simple measurement. Our detection system comprises a 1. generation CT scanner. To sample the complete set of projections required the object is translated and rotated at fixed tube - detector positions. We use a Philips MCN 165 x-ray tube as a radiation $ource. The detector system for normal energy integrated measurements is described elsewhere. Fig. 4 gives the results of a typical diffraction scan. It shows the images a phantom scan, which consists of a Lucite block with several holes filled with different sugar solutions. In addition to the normal transmission image one gets several scatter images, which show different contrasts. In the left upper corner of the scatter images particular sugar solution is evident which is completely invisible in the transmission image, owing to the fact that the total attenuation of this sugar solution and Lucite are identical. This result demonstrates that x-ray diffraction CT can provide useful additional information to a normal CT scan. The variation of image contrasts are in good agreement with energy-blurred scatter curves, Fig. 2b. 138 / SPIE Vol. 626 Medicine XIV/PACS IV (1986)

3 There are 3 sources of blurring the diffraction curves: 1. energy spread of the incident radiation; 2. angular width of the detector; 3. angular width of the object. The last problem can be solved using an iterative reconstruction technique. The second source of blurring is trivial and can avoided using small detector aperture. To show that the first problem can be solved for an x-ray source when energy sensitive detection is performed, we made energy resolved measurements with a Ge detector, which could be scanned by a stepping motor. The output of the detector was fed into a spectroscopy preamplifier (Ortec 472) and then into a multi-channel analyser (Canberra MCA 8100). The measured spectra of the scattered radiation were transferred into the memory of a LSI II/73 computer for further processing. To prove the feasibility of contrast enhancement, we measured the angular distribution of scattered radiation from a piece of Al Icm * Icm * Icm substending an angle of 0.1 deg. with a detector aperture of 1 mm. If one combines the good angular resolution with the energy-resolved processing one gets a very sharp scatter peak. Results are given in Fig. 5, which showes the gain in resolution if one transferms the energy resolved data into the corresponding pattern at a specific wanted energy. Here we choose 60 kev, which is the mean energy of our spectrum. At the moment we are not able to introduce this contrast enhancement into our imaging technique because one has to implement not only the energy resolution but one has increase the angular resolution to the same degree, otherwise a bad angular resolution would destroy all the gain obtained. A sufficient signal/noise ratio in combination with a short measuring time can be obtained by integration of the scattered radiation in annuli around the central beam, because the spatial distribution of the scattered radiation is circularly symetric around the primary beam for amorphous tissues. This requires a new detector design and will be a task for future activities. Conclusion We presented a new imaging method : X-ray Diffraction CT, which may offer useful additional information. Due the polychromatic radiation source the maximum obtainable contrast is reduced. We developed a solution for correcting this energy blurring by energy resolved photon counting. First results prove the feasibility of our proposed solution. References 1. Harding G., Kosanetzky J., Neitzel U., 1985, Phys. Med. Biol., 30, Harding G., Kosanetzky J., Neitzel U., 1985, Proc. XIV Int. Conf. Med. and Biol. Eng., Johns P.C., Yaffe M.J., 1983, Med. Phys., IO,40 4. Muntz E.P., Fewell T., Jennings R., Bernstein H., 1983, Med. Phys., 10, Neitzel U., Harding G., Kosanetzky J., 1985, accepted for publication in Phys. Med. Biol. primary beam, transtatton direction object Figure 1. Schematic drawing of a diffraction CT system SPIE Vol. 626 Medicine XIV/PACS IV (1986) / 139

4 i-i I I I 60 kev water _.-.r-- scatter 5O angle loo 120 kvp spectrum water... plex.... i polycarbo ;otysty ren O0 scatter 5O angle loo Figure 2. a) Scatter curves of different plastics for monochromatic x-radiation b) Scatter curves of the same plastics for polychromatic x-radiation 140 / SPIE Vol. 626 Medicine XIV/PACS IV (1986)

5 x -I-.- cn 5 C.- t /--- 8 \,* *- 0 \ /. \ 8 & total \ \ w.a c / O0 scatter angle Figure 3. Blurring of diffraction patterns by energy integration. The diffraction patterns for three different energies are presented together with the resultant total scatter curve Figure 4. Results of a diffraction CT scan: 1 Transmission and 9 scatter images at different scatter angles of a Lucite phantom with different sugar solutions are shown (scatter angles: 1.5, 1.7, 2.3, 4.1, 4.5, 5.0, 5.5, 6.1 ) 2.7, SHE Vol. 626 Medicine XIV/PACS IV (1986)

6 I! I I 1 I 1 I ,,+Q-,. c,+ci.+u.&& ~+++++t+,+,,.. I, I O0 5O loo scatter angle,' Figure 5. Scatter curves of a piece of Al (Icm * Icm * Icm) measured by a Ge detector with energy integration as well as with energy resolution (0 = energy integrated, + q energy resolved, processed for 60 kev) 142 / SPIE VoL 626 Medicine XIV/PACS IV (1986)