WO2013166574A1

WO2013166574A1 - Detection system for determining the location of scintillation events in a monolithic scintillation crystal

Info

Publication number: WO2013166574A1
Application number: PCT/BR2013/000158
Authority: WO
Inventors: Bonifácio Daniel Alexandre BAPTISTA; Maurício MORALLES
Original assignee: Comissão Nacional De Energia Nuclear (Cnen)
Priority date: 2012-05-08
Filing date: 2013-05-08
Publication date: 2013-11-14
Also published as: BR102012010830A2

Abstract

The present invention relates to a detection system for determining the location of scintillation events in a monolithic scintillation crystal. The system is formed by: a source of X-ray or gamma ray radiation, a scintillation crystal, an array of multiple photosensor elements coupled to one of the faces of the scintillation crystal, optical coupling means that provide an optical interface between the scintillation crystal and the photosensor matrix, a reflective optical coating for optically insulating the crystal and improving the optical photon detection efficiency of the photosensors, photosensor signal acquisition and processing in order to determine the location of the scintillation events; and data recording and image reconstruction.

Description

DETECTION SYSTEM FOR DETERMINING THE POSITION OF MONOLITHIC CRYSTAL SCINING WINDS "

TECHNICAL FIELD OF THE INVENTION

The present invention relates to a detection system for determining the position of monolithic scintillation crystal scintillation events. This system considers that when an X-ray or gamma-photon interacts in a scintillating crystal, the interaction site becomes a source of isotropic emission optical photons. The position of the interaction and the energy deposited by the X or gamma radiation photon can be determined by the radiation flux density of the scintillation optical photons measured at different positions in space. A process for acquiring and processing photosensor signals is then employed to determine the position of scintillation events. The present invention can be used in the field of Medical Physics - including in technologies such as Positron Emission Tomography (PET) and Single Photon Emission Tomography (SPECT) -, Nuclear Physics, Particle Physics, Astrophysics and any other application. for detecting ionizing radiation using scintillating crystals and requiring scintillation event location information. BACKGROUND OF THE INVENTION

The use of Nuclear and Particle Physics instrumental techniques in medical research, diagnosis and therapy applications has increased considerably in recent years. Certainly, medical imaging techniques have benefited most from technologies developed for particle detectors. This is evidenced by existing imaging modalities: Positron Emission Tomography (PET), Single Photon Emission Tomography (SPECT) Emission Computed Tomography "), Computed Tomography (CT), conventional X - ray and others.

The number of applications of the PET technique, which is based on the internal administration of radiopharmaceuticals, has been increasing in recent years, being used in the study of biochemical processes and blood flow of humans and other animals, such as the detection of tumors in very early stages. , which is not possible in other diagnostic imaging tests, such as computed tomography, magnetic resonance imaging and ultrasound. In addition, it is also applied in brain studies and in the development of new cancer drugs. Currently, PET scans have been associated with conventional x-ray tomography scans, allowing imaging of radiopharmaceutical distribution with great anatomical accuracy. This association is called PET / CT. Another hybrid technology is magnetic resonance PET (PET / MR), where the radiation doses absorbed by the patient are lower than the PET / CT technique.

There is a global effort to improve PET imaging systems, noted by the number of studies being conducted at various universities, research centers and companies in various countries. Several of these studies propose the development of dedicated PET scans for small regions such as small animals (ZIEMONS et al., 2005; CHATZIIOANNOU et al., 1999; YANG et al., 2004; ROLDAN et al., 2007; TAI et al. , 2005; SURTI et al., 2003; SPINELLIA et al., 2007; SCHÃFERS et al., 2005; GUERRA et al., 2006; ISHII et al., 2007; BALCERZYK et al., 2009; YAMAMOTO et al., 2010 ; YAMAYA et al., 2011), in order to study biochemical processes at molecular level for pharmacological, genetic and pathological investigations (GUERRA; BELCARI, 2007c, 2007a). Other applications are for parts of the human body, such as the brain (Séguinot et al., 2004), and studies and diagnosis of breast cancer (Motta et al., 2004; MOSES; Ql, 2004; NEVES, 201 1; WEINBERG et al.,

2005; MACDONALD et al., 2009; YANAGIDA et al., 2010), in one modality Positron Emission Lammography (PEM).

The purpose is to build more compact PET scanners with lower costs and even better imaging results, shorter examination time and lower amount of radiopharmaceutical administered for each specific application (MUEHLLEHNER; KARP, 2006) than a body PET scanner. would make it possible for these applications. The main advances in positron emission tomography for small structures are based on two aspects (WATANABE et al., 2002):

- Improved instrumentation to improve image quality parameters such as spatial resolution and detection system sensitivity, and the development of components to correct physical factors responsible for image degradation, such as gamma photon attenuation and events from Compton scattering.

- Improvement of the data acquisition process to obtain better image quality and more accurately quantify physiological parameters of clinical interest and scientific research.

Most PET systems developed today are based on the use of discrete scintillating crystals using Anger logic, where spatial resolution is influenced by the size of individual crystals. The use of smaller crystals improves spatial resolution but decreases sensitivity, as the higher granularity of the detector block scintillating crystals increases the volumes between crystals, which are not sensitive to X and gamma radiation. This also adds a higher cost to the system due to increased assembly complexity and the fabrication of crystals, which must be cut and polished using precision mechanics, and the fact that part of the crystal material is wasted during the same cut. .

In patent literature, Anger's logic is the subject of US301 1057 [Anger1961], referring to a new radiation detector and an instrument to more accurately measure the location of a distributed radiation source. Anger's logic determines the "center of gravity" of the relative intensities of the photomultiplier signals using an esistive chain. US7476864 [Bavciera2009] relates to a detector that determines the impact of gamma radiation position within scintillating crystal by calculating the Depth of Interaction (DOI) of radiation in the crystal. However, the Anger logic [Anger1961] used in monolithic scintillating blocks [Bavciera2009] is less accurate, as it presents slightly larger spatial distortions than the proposal of the present invention.

US2009 / 0242773 [Zhang2009] refers to the improved position estimation of photodetection events, where images were compared with Anger logic and substantially less image distortion was observed, which may allow automatic segmentation of the image. crystal from raw image data, which is particularly beneficial with regard to the simplification of the PET calibration system. However, this technique requires cutting the edges of the scintillating crystal. On the contrary, the present invention allows the use of scintillating blocks with larger volume than that presented in [Zhang2009], there is no need of cutting and, consequently, avoiding the waste of crystal material and reducing the manufacturing costs of the crystals (which must be cut using precision mechanics).

Patent application US2010 / 0044571 [Miyaoka2010] deals with a method for determining the three-dimensional position of a scintillation event. It is also necessary to calibrate the equipment using "Look Up Tables" (LUT) tables. obtained experimentally for each detector. On the other hand, in the present invention the only calibration required is energy calibration.

US6459085 [Chang2002] and US7737407 [Grazioso2010] refer to a system for determining DOI in nuclear imaging and a method and device for providing DOI using the PET technique, respectively. Both techniques make use of light guides. In contrast, the present invention does not use said guides, that is, the The detectors are coupled directly to the scintillating crystal. This facilitates detector construction, and increases the detection efficiency of optical photons, which improves detector energy resolution.

US7956331 [Lewellen201 1] and US7482593 [Shao2009] also refer to a method of determination or device for determining a scintillation event using discrete scintillating crystals, unlike the present invention using monolithic scintillating crystals. Again, the use of monolithic crystals reduces the cost of crystal manufacturing and the waste of crystal material due to precision cutting of the edges.

Among non-patent references, the work of [LI2010] LI, Z. et al. Nonlinear least-squares modeling of 3d interaction position in a monolithic scintillator block. Physics in Medicine and Biology, v. 55, no. 21, p. 6515, 2010, is another method that does not use LUT, but determines the position of the flicker event using the individual signals from the photosensor array. In the present invention calculations are made by summing the row and column signals, causing the number of channels to be reduced from rxs to r + s.

The following works also represent prior art references. They allow the use of discrete scintillating crystals, unlike the present invention, which uses monolithic scintillating crystals:

[Huber2003] HUBER, J. et al. Development of the LBNL positron emission mammography camera. IEEE Transactions on Nuclear Science, v. 50, no. 5, p. 1650-1653, 2003.

[Surti2003] SURTI, S. et al. Design evaluation of A-PET: A high sensitivity animal PET camera. IEEE Transactions on Nuclear Science, v. 50, no. 5, p.

1357-1363, 2003.

[Abreu2006] ABREU, M. et al. Design and evaluation of the clear-pem scanner for positron emission mammography. IEEE Transactions on Nuclear Science, v. 53, no. 1, p. 71 - 77, 2006.

[Raylman2006] RAYLMAN, RR et al. Development of a dedicated positron emission tomography system for the detection and biopsy of breast cancer. Nuclear Instruments and Methods in Physics Research Section A: Accelerators, Spectrometers, Detectors and Associated Equipment, v. 569, no. 2, p. 291-295, 2006.

[Yang2006] YANG, Y. et al. Depth of interaction resolution measurements for a high resolution PET detector using position sensitive avalanche photodiodes. Physics in Medicine and Biology, v. 51, p. 2131-2142, May 2006.

The articles below are only bibliographic references to better understand the concepts adopted in the invention:

[America1995] AMERICA, O. S. of. Fundamentals, Techniques, and Design. 2. ed. New York: McGraw-Hill, 1995. 42.9-42.13 p.

[Hecht2002] HECHT, E. Optics. 4. ed. San Francisco: Addison Wesley, 2002. Cap 4 and Cap 3.

[Spiegel1992] SPIEGEL, M.R. Theory and Problems of Probability and Statistics. 2nd ed. New York: McGraw-Hill, 1992. p. 1 14-115.

It is therefore concluded that the present invention has unique features and advantages over the state of the art. That is, the detection system proposed here uses monolithic scintillating crystals, avoiding the disadvantages of using discrete scintillating crystals (high crystal manufacturing cost due to cutting and polishing and material waste) associated with a process of acquisition and processing. the most accurate and least non-linearly distorted photosensor signals aimed at obtaining better image quality parameters, reducing the examination time and quantifying the radiopharmaceutical dose administered for each specific application. In choosing the detector block, consideration was given to using as few photodetectors / acquisition channels as possible, reducing system cost while maintaining system spatial resolution at or better than that value. PET systems available today. In this sense, the present system uses the sum of the row and column channels, which reduces the number of photodetectors / acquisition channels from rxs to r + s. Another important feature is the Depth Of Interaction (DOI) wind inside the sparkling crystal. The DOI information significantly improves spatial resolution throughout the system's field of view, as it better determines the position of the X-ray or gamma photon interaction, avoiding parallax errors. DESCRIPTION OF THE INVENTION

The description of the invention refers to the following figures, in which: Figure 1 illustrates the side view diagram of the detection system representation exemplifying the generation of the scintillation event by interaction of an X-ray or gamma photon on the scintillating crystal. and the subsequent acquisition and processing of photosensor signals.

Figure 2 shows the side view diagram of the detection system representation for a front reading.

Figure 3 illustrates the side view diagram of the detection system representation for a rear reading.

Figure 4 shows the diagram representing the inline and column summation reading of the photosensor signals from the 8 x 8 photosensor matrix.

Figure 5 illustrates the scintillation event occurring at position (xo, y ₀ , z ₀ ) after energy deposition by a gamma or X-ray photon and the detection of scintillation-produced optical photons by an array element of photosensors positioned at (x _p , y _q ).

Figure 6 shows the flowchart of the process of determining the position of the 2D interaction of the X or gamma radiation photon.

Figure 7 illustrates three virtual sources representing the reflections on the crystal faces, which have a reflective material coating. The positions of the virtual sources are relative to the position of the actual source.

Figure 8 represents the set of optical layers in the detector block and their respective thickness and refractive index values for a wavelength of 420 nm, which corresponds to the peak emission wavelength of an LYSO scintillator. Figure 9 shows the adjusted transmittance data calculated for a beam and optical photons traversing the epoxy and Meltmount resin layers as a function of the initial beam angle of incidence.

Figure 10 presents the general scheme of the detection system used for positron emission tomography.

The present invention relates to a detection system for determining the position of monolithic scintillation crystal scintillation events. This system considers that when an X or gamma radiation photon interacts within the scintillating crystal, the interaction site becomes a source of isotropic emission optical photons. The position of the interaction and the energy deposited by the X or gamma radiation photon are determined by measuring the radiation flux density of the scintillation optical photons in different positions of space. Figure 1 shows an exemplary diagram of said detection system in which a scintillation event (1) is generated by means of an X or gamma radiation photon (2) which hits the input surface (3) and undergoes an interaction within of the scintillating crystal (4). Said system consists of X or gamma radiation (2), a scintillating crystal (4), a set of multiple photosensor elements (5) coupled to one of the faces of (4). The detection system further includes other components which are described below. Optical coupling means (6), such as a gel or an optical grease or an epoxy based compound or other type of adhesive resin, such as Meltmount resin, not limited to these, is used to provide an optical interface (or coupling) between (4) and (5). A reflective optical coating (7) such as Teflon or 3M Enhanced Specular Reflector (ESR), not limited to these, is used to optically isolate the crystal and improve the detection efficiency of optical photons by photosensors. A process (8) of acquiring and processing signals (5) is performed to determine the position of scintillation events (1).

The scintillating crystal is typically and / or approximately a rectangular solid. Some possible scintillating crystals are: lutetium oxyortosilicate cerium doped opium (LSO), lutetium fine silicate (LFS) and cerium-doped yttrium yttrium oxyortosilicate (LYSO), not limited to these.

Photosensors can be of various types such as, for example, photomultiplier tubes (PMT), silicon photodiodes, avalanche photodiodes (APD), and silicon photomultiplier (SiPM), not limited to them. These multiple photosensor elements represent a two-dimensional "photosensor matrix" (2D).

Figure 2 illustrates, as one embodiment of the invention, the Front Side Readout (FSR) scheme in which X-ray or gamma-2 photons fall into (4) from the input surface (3). ), which is the face on which the photosensor array (5) is coupled. In this embodiment, (5) must have a low attenuation of (2) for the use of a front reading. Another embodiment of the invention, Figure 3, shows a diagram of the Back Side Readout (BSR), which is where (2) focuses on (3), which is the opposite face on which the photosensor array (5) is coupled. Said detection system of this invention may perform front, back and side reading.

Figure 4 shows a reading diagram of the (5) signals in row (9) and column (10) sums representing a further embodiment of the invention specifically for an 8 x 8 photosensor array array.

The process of acquisition and processing of photosensor signals determines the position of gamma or X-ray photon interaction within the scintillating crystal. Such a process may be implemented via hardware and / or software. Figure 5 shows the conventions defined for the x, y and z axes in a monolithic scintillation crystal volume (4) and the generation of a scintillation event (1) at position (x ₀ , yo, Zo) by the deposition of energy of a gamma or X-ray photon (2). The xy plane defines the plane of the position-sensitive photosensor array (5), optically coupled to (4), and the z ₀ position is the interaction depth of (2) within (4). The position (x _p , y _q ) represents the center of one of the matrix elements of (5), where the optical photon detection. In this example, the matrix of (5) is coupled to the face of the input of (2) (front reading).

1. 2D position determination

A possible process for determining the values of the x ₀ and yo coordinates of gamma or X-ray photon interaction is based on the weighted average calculation with trimming and removal of outliers, that is, points that deviate from the distribution trend. signals collected.

Consider an array of photosensors of rxs elements. Let lj be the intensity of channel j, given by the sum of the signal intensities along the u axis of the photosensor matrix, where u can be the x or y axis. The index j ranges from 0 to n-1, where n is the number of channels along u, that is, n can be r or s. Given all values of the intensities 1j of each channel and their respective positions Uj along the u axis, the process for determining the position of the interaction u ₀ :

1. Determination of the highest intensity channel l _max and its respective index j _max .

2. Select all channels with intensity where and

0, 1 <F <0.8 Scan of j _max index channels up to 0 and discard all intensities channels

Scan jmax index channels to n-1 and discard all channels with intensities

3. New selection of index channels if only channel

of greater intensity has not been discarded

4. Calculation of the uo position by equation 1.1 as the average position of the selected channels, weighted by the selected intensities subtracted from the value

The factor F has a typical value of 0.4 and depends on the crystal thickness and taterial. Factor F should be chosen so that position determination has an optimal result.

Said process is graphically represented in Figure 6.

The method may or may not be used in conjunction with other methods for determining any of the Xo, yo, and ZQ coordinates.

2. 3D position determination

Another possible process that can be used to determine the Xo, yo, and z ₀ coordinate values of the gamma-ray or X-ray interaction adopts a distribution model of the incident photons in the photosensor matrix. Figure 7 shows the convention adopted by the optical photon distribution model incident on the photosensor matrix (5) to define three virtual sources to describe the reflections on the crystal faces, which have a reflective material coating and their positions are defined as follows. according to the position of the actual source.

The implemented optical photon distribution model is expressed by two functions, Nphj and Nph _j , which describe the number of optical photons collected from the sum of the signals of row i (x axis) and column j (y axis), respectively, of the matrix. of photosensors:

where Nph _d (Xj; xo; z ₀ ) and Nphd (yj; yo; z ₀ ) are the sums of the line and column signals, respectively, of the optical photons that are collected and that directly affect the plane of the photosensor matrix. .

The sum of each of the functions describes the three virtual sources, positioned at x ₀ + Xk, yo + yk and z ₀ ⁺ Zk, which represent the contributions of specular reflections at the reflecting / scintillating crystal interface. All collected optical photons that are not described by the above terms are considered as background radiation. Such photons have suffered at least one They do not reflect the spacial position of the gamma or X-ray photon, but their importance lies in the fact that they contribute to improving the energy resolution of the collected signal. The projection distribution in one dimension of these photons is considered uniform, being represented by the constant NpheG-2.1. Cauchy-Lorentz distribution

The inverse square law of distance does not apply when the dimensions of the detectors are not negligible when compared to the distance from the light source. Therefore, the optical photon distribution model assumes that the projection on the xz and yz planes of the distribution of the optical photons emitted by the energy deposited by an X or gamma radiation photon at the position (xo; yo; zo) and that directly affects the xy plane of the photosensor matrix follows the continuous probability distribution of Cauch-Lorentz [SPIEGEL1992]:

where u is the x or y position along the projection axis relative to the intensity to be determined, uo is the xo or yo position of the radiation photon or gamma interaction on the projection axis and z ₀ is the position of the interaction depth inside the crystal. This distribution is normalized by the area with the following factor:

2.2. Optical photon transmission through photodetector window and optical resin

In the model was implemented the effect of optical photon transmission through the photosensor window and the optical resin, which is described by the transmittance of an optical photon beam as a function of the angle of incidence.

The reflectance and transmittance of an optical photon beam on a two-media interface with distinct refractive indices is described by the Fresnel equations [HECHT2002a]. However, the detector block used in this As an example, it has three interfaces (crystal / resin, resin / epoxy and jpoxy / photosensor), so multiple reflections may occur. These reflections can interfere constructively or destructively, and Fresnel's equations do not take this phenomenon into account. Thus, the transfer matrix method [AMERICA1995] was used, which is a method employed to analyze the propagation of electromagnetic waves across a series of layers with different thicknesses and refractive indices. The method is based on a matrix formulation of thin film boundary conditions using Maxwell's equations [HECHT2002b]. For an L-layer system, the method consists of calculating the following matrix product:

being

Where

parallel polarization (2.8) or

p ₀ | _ar j _zacã perpendicular (2.9) The angle is obtained by the angle of incidence θο using Snell's law:

where n is the refractive index of the incident medium, that is, of the scintillating crystal. Thus, the electric vector E ₀ and the magnetic vector Ho can be calculated:

nde η ₃ is the effective refractive index of the substrate, which corresponds to silicon in the housing of a silicon diode photosensor. The transmittance is defined by:

Since the scintillation light is non-polarized, the transmittance shall be calculated for parabolic (T _p ) and perpendicular (T _s ) polarization, and the simple mean of the two values shall be determined:

In addition, we must make a weighted average of the transmittances obtained according to the wavelength intensities of the crystal emission spectrum:

Figure 8 shows the layer set of the detector block optics chosen as an example, where a beam of light within the scintillating crystal falls at angle θ ₀ at the interface with the Meltmount optical resin and can be reflected at angle Θ _Γ or refracted with angle θ ₂ , and so on. The refractive index and thickness values of each layer are also indicated.

Thus, the matrix M was determined to calculate the transmittance of the detector block optical system:

With the transmittance data, an exponential function has been adjusted that roughly describes this behavior for incidence angles smaller than the critical angle:

This same function can be represented as a function of the position of the interaction of the X or gamma radiation (u ₀ ; z ₀ ) and the position u of the intensity of the optical photons collected by the matrix:

where u is the x or y coordinate along the projection axis relative to the intensity to be determined, Uo is the Xo or y ₀ coordinate of the radiation photon or gamma interaction on the projection axis and z ₀ is the position of the depth of interaction within the crystal. The adjustment was made using the nonlinear least squares method (Levenberg-Marquardt algorithm) and the parameters found were:

Table 1. Values of the fit parameters of equation 2.17 using the nonlinear least squares method.

Figure 9 shows a graphical representation of the calculated transmittance data and the exponential function setting, valid from 0 to 90 degrees.

As can be seen from the graph, the function has been adjusted from zero to the critical angle, which is approximately 56.85 degrees. Above the critical angle, the function provides negative transmittance values, which have no physical significance. Therefore, the model considers these negative values to be zero.

2.3. Full Signal Probability Distribution

Taking into consideration all the factors described above, the

Probability distribution of the signal of the collected optical photons that directly affect the plane of the photosensor matrix is:

where A _d is the distribution normalization factor and Au is the photosensor dimension along the corresponding axis. The integral makes the model take into account the photosensor dimensions, which are not negligible when compared to the possible distances from the position of interaction to the photosensor matrix. Since equation 2.17 is valid only for incidence angles greater than or equal to zero> 0, the same condition applies for equation 2.18. However, the result can be determined to negative angles because of the existing symmetry. Solving the definite integral:

we have:

Defining the response function R of the photosensor array element I as:

we have:

Similarly, the probability distributions of each of the three virtual sources are given by:

where A _v1 , A _v2 and A _V 3 are the normalization factors for each of

distributions. Normalization factors are calculated through the following relationships:

where Nphtotai is the total number of optical photons collected by the photosensor array, N _and i is the number of elements of the photosensor array, and N _c is the number of channels for the u axis.

2.4. Model Parameter Adjustment

The model parameters were adjusted using two function adjustment methods: the maximum likelihood method and the nonlinear least squares method. However, other adjustment methods may be used.

2.4.1. Limits and initial values of model parameters

For the adjustment of the signal distribution model data to be performed correctly, it is important that the parameters to be adjusted have initial values close to the actual values, in order to prevent the process of enmity from converging to the wrong minimum point. In addition, the parameters are limited to a range of values for physical significance and to avoid divergence of the estimated values.

2.4.1.1. Background Radiation (Nph _B G)

The initial value of the parameter representing the background radiation (Nph _B G) is defined as the simple mean of the two plus signs.

low intensity collected by the sum in row or column by the photosensor matrix:

An upper limit equivalent to the initial value has been set for this parameter. The lower limit has a typical value of 60% of Nph _B G, but can vary from 0 to 90%.

2.4.1.2. Interaction Position (xo and yo)

One process that can be used for the initial estimation of the interaction position in the xy plane (x ₀ and yo) is that of item 1 (2D Position Determination). However, other processes can be used [Malmin2010],

The upper and lower limits of the initial value u, represented respectively by u _sup and Ujnf, were defined with the following values and conditions:

.4.1.3. Interaction Depth (zo)

> initial interaction depth value z ₀ and its upper limit is defined as the crystal thickness value. The lower limit is set to zero.

2.4.2. Adjustment Methods

2.4.2.1. Maximum likelihood method

The Maximum Likelihood (ML) method is a well-known statistical method used to estimate the parameters of a statistical model from a set of data to be adjusted. Given a set of N-detected optical photons. inline / column channels {m ,; i = 1,2; ...; Nc}, where Nc, we have that the probability that mi is obtained in the range of rri2 is obtained in the range and

so successively, is given by:

Where

^ the value of the probability density function for mi, where true parameters are defined by the vector

. Therefore:

Since the factor is constant, the same can be

neglected for the definition of the likelihood function L, which is given by:

Assuming that the values of m, obey a Poisson distribution, the function L is given by:

where λ, = Nphj.

The maximum likelihood method is to estimate the true parameters

through a vector that maximizes the function of Since the L function is positive, it is more convenient to maximize logarithm of the L function, which has the same maximum point:

where ln (mj!) is a constant. Thus, the estimate of the true value

is determined with the following estimator:

2.4.2.2. Minimum squares method

The least squares method is one of the most widely used techniques for determining the best fit of a model or function to an uncorrelated data set. The method uses the concept of residue, which is the difference between an observed value m and its respective adjusted value Nph (u ,; P):

The method consists in maximizing the fit with an estimator that minimizes the sum of the squares of the regression residuals. The least squares method estimator is given by:

In the least squares method, the number of observed data "N" must be greater than the number of parameters to be adjusted. The method is divided into two categories: linear, used when the parameters have a linear relationship to each other, and nonlinear, used in nonlinear problems where the solution is obtained by iterative methods. When data uncertainties differ, a weighted estimator approach should be used:

ide σ, is the uncertainty of the value nrij.

Positron Emission Tomography

One of the non-limiting embodiments of the invention is shown in Figure 10, which relates to a detection system employed in the Positron Emission Tomography (PET) technique. The PET technique is a noninvasive nuclear medicine imaging method used to obtain the distribution map of a radioactive tracer, called a radiopharmaceutical (1), internally administered to a living organism (12). In general, (11) is a biologically relevant chemical labeled with a short half-life positron-emitting radionuclide such as carbon-1, nitrogen-13, oxygen-15 and fluorine-18. Gamma radiation photons (2) from positron annihilation interact with the scintillating crystal (4). The collected signals go through a signal acquisition and processing unit (13). Recorded data and image reconstruction are performed using a computer (14).

Claims

1. Detection system for determining the position of scintillation events in monolithic scintillating crystal, characterized in that it consists of:

(a) a source of radiation photons;

(b) a monolithic scintillating crystal;

(c) an array of multiple photosensor elements coupled to one side of said scintillating crystal;

(d) An gel or optical grease or epoxy based compound or other type of adhesive resin to provide optical interface between the scintillating crystal and the multiple photosensor assembly;

(e) a reflective optical coating;

(f) A process for acquiring and processing photosensor signals;

(g) A process of recording data (signals) and reconstructing the image.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the source of radiation photons can emit gamma rays or X-rays.

Detection system for determining the position of monolithic scintillating crystal scintillation events according to claim 1, characterized in that the scintillating crystal comprises cerium-doped lutetium oxyortosilicate (LSO), lutetium fine silicate (LFS) ) and cerium-doped yttrium yttrium oxyortosilicate (LYSO).

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the multi-element array Photosensors consists of an array of two-dimensional photosensors.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 4, characterized in that the photosensors comprise photomultiplier tubes (PMT), silicon photodiodes, avalanche photodiodes (APD) and silicon photomultipliers (SiPM).

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the optical resin or gel comprises Epoxy or Meltmount resin.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the reflective optical coating consists of an insulating material used to optically isolate the crystal and improve the efficiency of detection of optical photons by photosensors.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 7, characterized in that the insulating material comprises Teflon and 3M ESR ("Enhanced Specular Reflector").

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the process of acquisition and processing of the photosensor signals consists of the following procedures:

(a) Collection of radiation photons which affect the input surface of the scintillating crystal;

(b) Reading the signals generated by the photosensor matrix; (c) Determination of the position of the interaction of optical photons within the scintillating crystal and the deposited energy.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 9, characterized in that the reading of the signals generated by the photosensor array comprises the front, side or rear reading according to Photosensor array is coupled to one side of the scintillating crystal.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 9, characterized in that the reading of the signals generated by the photosensor matrix consists of the sum of the inline and column signals (rxs elements).

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 9, characterized in that the determination of the position of the interaction of the optical photons within the scintillating crystal and the deposited energy consists in measuring the radiation flux density of said photons at different positions of space.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 9, characterized in that the determination of the position of the interaction of optical photons within the scintillation crystal comprises the two-dimensional planes of the photosensors and three-dimensional.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 13, characterized in that the determination of the two-dimensional position of the photosensor matrix consists of calculating the weighted average with trimming and removal of points that deviate from the distribution trend of the collected signals.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 13, characterized in that the determination of the three-dimensional position consists of an estimate of parameters of a signal distribution model collected by the photosensor matrix elements.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 15, characterized in that the distribution pattern of the signals collected by the photosensor matrix elements consists of calculating the transmittance of each light beam, as a function of its angle of incidence, generated by the optical photons that traverse the photosensor array and the optical coupler medium.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 15, characterized in that the distribution pattern of the signals collected by the photosensor array elements consists of calculating the intensity of the photons optics, as a function of the position of the interaction, which directly affect the photosensor matrix.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 15, characterized in that the distribution model of the signals collected by the photosensor matrix elements consists in calculating the contribution of the reflections at the optical coating / scintillating crystal interface for the signal strength collected by the photosensor matrix.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 15, characterized in that the determination of the initial position value of the three-dimensional interaction comprises the process of determining the two-dimensional position.

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that the signal acquisition and processing process comprises the use of software or analog electronics techniques. .

Detection system for determining the position of monolithic scintillation crystal scintillation events according to claim 1, characterized in that its functional architecture can be employed in imaging techniques, comprising positron emission tomography (PET) and single photon emission computed tomography (SPECT).