Method and system for three-dimensional imaging of transparent biological objects within biological specimens by full-field optical tomography

Abstract

The present specification relates to a three-dimensional imaging system (100), comprising a light source (110) configured to illuminate a biological sample (10) through which it passes and configured to emit a beam of spatially incoherent light having a given central length, an optical imaging system (120) comprising a microscope lens (121) with a given object focal plane (125) in the vicinity of which the sample (10) is positioned, a means for axially moving the microscope lens relative to the sample, a two-dimensional acquisition device (140) optically conjugated with the object focal plane, comprising a number of elementary detectors arranged in a detection plane (141), and a processing unit (150). For each slice of the biological object of the sample, a number of two-dimensional interference signals resulting from the optical interference between the illumination beam and the beam scattered by the field of the slice are acquired, and at least a first image is calculated from the number of two-dimensional interference signals.

Description

[Technical field]

[0001] The present disclosure relates to a method and system for three-dimensional imaging of transparent biological objects within biological samples by full-field optical tomography, which is particularly applicable to cellular and subcellular imaging. [Background technology]

[0002] Although biological objects, especially cells or clusters of cells, can be examined at the submicron scale using optical microscopy, most biological research uses fluorescence microscopy, which requires genetic or chemical modification of the sample.

[0003] In cases where sample preservation is essential, a few alternatives based on the intrinsic optical properties of biological objects have been developed, among which phase contrast microscopy (see, for example, Non-Patent Document 1), differential interference contrast microscopy (see, for example, Non-Patent Document 2) or digital holography (see, for example, Non-Patent Document 3) allow the visualization of cells in 2D and 3D environments.

[0004] Optical tomographic imaging techniques make it possible to generate deep cross-sectional images of biological objects. Among these techniques, optical coherence tomography (OCT), which is based on spectral broadband optical interference microscopy, and confocal microscopy using geometric filtering, are particularly well known. Both techniques require a two-dimensional scan of the sample to generate an image of its front surface.

[0005] More recently, techniques have been developed for acquiring images by non-coherent light full-field interferometric microscopy, known as full-field OCT (or FF-OCT). Full-field OCT imaging techniques are described, for example, in Non-Patent Document 4 and in Patent Document 1.

[0006] The full-field OCT imaging technique is based on the use of light backscattered by the sample when illuminated by a light source with low temporal coherence length, and, especially in the case of biological samples, by microscopic cellular and tissue structures. This technique exploits the low temporal and spatial coherence of the light source to isolate the light backscattered by a virtual deep slice in the sample. The use of an interferometer allows, by interference phenomena, to selectively generate an interference signal representative of the light coming from a given slice of the sample, and to eliminate the light coming from the rest of the sample.

[0007] Full-field OCT imaging techniques make it possible to obtain three-dimensional images with a typical resolution of the order of 1 μm, which is higher than the resolution of the order of 10 μm that is often obtained with other conventional OCT techniques, such as spectral domain OCT (known by the acronym "Fourier domain OCT"). With such a high resolution it is possible to visualize most of the tissue structures, in particular blood vessels, their walls, collagen, fat cells, etc. In addition, this technique is particularly fast: thus, with a full-field OCT microscope, images of a few centimeters can be obtained. 2 It is now possible to generate images representing deep slices with surface areas of up to 100 mm in just a few minutes.

[0008] In addition to obtaining a structural map of a biological sample, another challenge for intrinsic optical imaging is the extraction of specific information that deduces the biochemical properties at the cellular level.

[0009] US Patent No. 5,399,633 describes an imaging method by full-field interference microscopy called DC-FFOCT ("Dynamic Contrast FFOCT"), which allows access to information not perceptible with images obtained according to conventional full-field OCT techniques, such as the internal structure of cells (in particular membranes, nuclei, cytoplasm). This imaging method is based on the analysis of the temporal variations in the intensity of the interference signal induced by the variations in the path difference associated with the movement of, for example, vesicles, mitochondria, intracellular organelles. Such a dynamic analysis makes it possible to obtain full-field tomographic images with excellent contrast in fresh tissues in which the cells are still alive, and thus to visualize structures of biological subjects that were not perceptible with OCT before, since the backscattering of light by cells is very slight compared to the structures of tissues, such as collagen.

[0010] However, the FF-OCT and DC-FFOCT techniques described above work in reflection and are not compatible with highly transparent biological objects placed on transparent supports such as, for example, clear-bottom glasses, or plastic Petri dishes, or glass slides, etc. Indeed, the signal reflected by the support is usually 3-5 orders of magnitude larger than the signal backscattered by the biological object, e.g., cells, making it difficult to remove this signal. [Prior art documents] [Patent documents]

[0011] [Patent Document 1] French Patent Invention No. FR2817030 [Patent Document 2] French Patent Invention No. FR3034858 [Non-patent literature]

[0012] [Non-Patent Document 1] Lacey, A., Phase contrast, Light Microscopy in Biology: A Practical Approach, edited by Lacey, A., Oxford University Press, Oxford, UK, p23-25, (1999) [Non-Patent Document 2] Pluta, M., "Nomarski's DIC microscopy: A review", Proceedings of SPIE 1846:10-25, (1994) [Non-Patent Document 3] Amos, WB and White, JG. "How the confocal laser scanning microscope entered biological research." Biology of the Cell 95:335-342, (2003) [Non-Patent Document 4] Harms, F, Latrive, A and Boccara, AC, "Time Domain Full Field Optical Coherence Tomography Microscopy", Wolfgang Drexler book, Optical Coherence Tomography, page 791, editor: James G. Fujimoto, Springer, 2012. Summary of the Invention [Problem to be solved by the invention]

[0013] The present specification describes a new method for 3D imaging of biological samples by full-field optical tomography adapted for 3D imaging of transparent biological objects such as cell cultures or highly transparent thin structures with excellent contrast and ease of implementation. [Means for solving the problem]

[0014] As used herein, the term "comprising" is synonymous with "including" and "containing" and is inclusive or open ended and does not exclude other elements not described or expressed. Furthermore, as used herein, the terms "about" and "substantially" are synonymous with (meaning the same thing) a lower and / or upper margin of 20% of the respective value, e.g., 10%.

[0015] The present specification relates according to a first aspect to a method for three-dimensional imaging of a transparent biological object in a biological sample by full-field optical tomography, said three-dimensional imaging method comprising: - positioning the sample in the vicinity of an object focal plane of a microscope lens that encompasses a given optical axis; - illuminating said sample during transmission with an illumination beam of spatially non-coherent light having a given central wavelength; - displacing the microscope lens relative to the sample along an axial direction parallel to the optical axis of the microscope lens to define a plurality of positions of the sample relative to the object focal plane of the microscope lens, each position corresponding to a slice of the transparent biological object centered on the object focal plane of the microscope lens; - generating at least one first image of a field of the section for each location of the sample; generating at least one first image of the scene of the slice; - acquiring a plurality of two-dimensional interference signals resulting from optical interference between the illumination beam incident on the field and a beam scattered by the field using a two-dimensional acquisition device comprising a plurality of elementary detectors arranged in a detection plane, the detection plane being optically conjugated with the object focal plane of the microscope lens by an imaging optical system comprising the microscope lens; - calculating the at least one first image from the plurality of two-dimensional interference signals using a processing unit; Includes:

[0016] In the present specification, a transparent biological object is understood to be any biological object with low scattering (usually less than about 5% of the incident light scattered by the object) through which the light from the illumination beam is likely to pass, such as a cell or a plurality of cells, such as a cluster of cells, a cell scaffold, a cell mat or biofilm, a thin slice of tissue such as prepared in anatomical pathology, etc. Typically, a transparent biological object for which an image is sought to be (at least partially) formed using the method according to the present specification has microscopic dimensions (between about 1 μm and about 2 mm). More precisely, when it is a cell or a plurality of cells, such a transparent biological object has dimensions between about 2 microns and about 50 microns.

[0017] According to the present specification, the transparent biological object forms part of a biological sample, i.e. a support on which the transparent biological object is placed, such as, but not limited to, a 2D or 3D cell culture in a transparent polymer, or the product of a fine needle biopsy (FNA) or "fine needle cytology", i.e. the product of cell and tissue sampling using a fine needle and a syringe.

[0018] In the 3D imaging method object herein, a variation in the phase of the scattered beam near the object focal plane of the microscope lens, called Gouy phase, is used to introduce a phase shift between the two-dimensional interference signals of the plurality of two-dimensional interference signals acquired to form the at least one first image. It is possible to show that beyond a certain distance from the object focal plane of the microscope lens, there is no phase shift between the interference signals. Thus, by displacing the object focal plane of the microscope lens within a slice of a given thickness of the biological object to be imaged, or by using the intrinsic movement of the internal structure of the biological object within the same slice of this given thickness, it is possible to obtain two-dimensional interference signals with different phase shifts from which an image can be calculated.

[0019] The image of the biological object formed using the imaging method according to the present specification is therefore an image of a field in a slice of a given thickness, referred to herein by the term "section" or "optical slice", hence "optical tomography". Applicants have shown that the thickness of this slice is of the order of the depth of field of the microscope lens and therefore depends on the numerical aperture of the microscope lens and the central wavelength of the illumination beam.

[0020] Thus, in the method according to the present invention, the field of the slice of the biological object on which an image is to be formed can be defined by a volume of the biological object consisting of a set of volume elements or "voxels", each voxel corresponding to a length equal to the depth of field of the microscope lens and to a cylindrical volume of the slice defined by the diffraction disk of the microscope lens. The lateral dimension of the field (in a plane perpendicular to the optical axis of the microscope lens) is equal to the lateral dimension of the image field defined by the effective detection surface of the detection plane divided by the magnification value of the imaging optical system including the microscope lens. In an exemplary embodiment, the effective detection surface encompasses the entire detection surface on which the elementary detectors or "pixels" of the two-dimensional acquisition device are arranged. In another exemplary embodiment, the effective detection surface can be limited to a region of the detection surface, for example by using a field stop. In practice, a voxel or volume element of the field of view of the slice in the object space of the microscope lens can correspond in image space (detection plane) to a number of elementary detectors, also referred to in this specification as "pixels", of the two-dimensional acquisition device in order to meet a sampling criterion.

[0021] According to one or more exemplary embodiments, when dealing with central wavelengths of the illumination beam in the visible range (400 nm-700 nm), the thickness of the slice of the object imaged using the method according to the present disclosure is between about 200 nm and about 10 microns for numerical apertures of the microscope lens between 1.25 and 0.3. In practice, for a given numerical aperture of the microscope lens, it will be possible to work with shorter wavelengths for imaging thinner slices of the object. The resolution of the image of the field within the slice is defined by the maximum lateral dimension of a voxel. Within the visible range, and for numerical apertures of the microscope lens between 1.25 and 0.3, the resolution is between about 0.25 microns and about 1 micron.

[0022] Relative displacement of the microscope lens with respect to the sample along an axial direction parallel to the optical axis of the microscope lens therefore makes it possible to create images of multiple sections of the transparent biological object to create a three-dimensional image of the object.

[0023] The three-dimensional imaging method described thus enables tomographic imaging, including of transparent biological objects, using a novel optical arrangement in which the sample is illuminated during transmission.

[0024] According to one or more exemplary embodiments, the multiple two-dimensional interference signals are obtained for different positions of the object focal plane within the thickness of the slice, resulting in multiple predetermined phase shifts between the illumination beam and the scattered beam ranging between -π / 2 and π / 2.

[0025] In the first embodiment described above, the method therefore comprises displacing the microscope lens relative to the sample along an axial direction parallel to the optical axis of the microscope lens to create an image of a field of a slice of the biological subject, and using the two-dimensional interference signal to calculate the image of the field of the slice being acquired for various positions of the subject focal plane of the microscope lens within the slice, thus enabling the two-dimensional interference signal to be acquired with a predetermined phase shift.

[0026] According to one or more exemplary embodiments, the calculation of the at least one first image comprises a linear combination of two-dimensional interference signals of the plurality of two-dimensional interference signals of a plurality of signals, for example a subtraction of two of the two-dimensional interference signals.

[0027] In practice, the calculation of the at least one first image comprises calculating, for each elementary detector of the two-dimensional acquisition device, a pixel value as a function of a linear combination of the intensities of the two-dimensional interference signals acquired by the elementary detectors.

[0028] In this specification, a "pixel value" is understood to be the value of the signal measured by the corresponding pixel or elementary detector.

[0029] According to one or more exemplary embodiments, calculating a pixel value is a function of a linear combination of the intensities of two-dimensional interference signals obtained by varying the distance between the microscope lens and the sample between about 1 / 10 of the depth of field and the depth of field, advantageously between about 1 / 10 and about 1 / 2 of the depth of field.

[0030] According to one or more exemplary embodiments, the relative displacement of the microscope lens with respect to the sample follows a periodic function of maximum amplitude λ / 4, where λ is the central wavelength of the illumination beam, such as a continuous periodic function, such as a sinusoidal function, or a periodic gradient, with a period determined as a function of the acquisition rate of the 2D detector to allow a temporal sampling adapted to the calculation of the image, such as two, three or four acquisitions per period.

[0031] According to one or more exemplary embodiments of the three-dimensional imaging method according to the first aspect: - the two-dimensional interference signal of the plurality of two-dimensional interference signals is acquired for a fixed position of the microscope lens relative to the sample, and calculating the at least one first image of the field of the section; - calculating, for each elementary detector of said two-dimensional acquisition device, at least one pixel value as a function of the value of a parameter representative of the temporal variation in the intensity of said two-dimensional interference signal acquired by said elementary detector.

[0032] In this second embodiment, the image of the field of the section of the transparent biological object is created without displacement of the microscope lens relative to the sample. The image differs from the calculation in the first embodiment because the movement of the intrinsic structure of the biological object and therefore the environment is characteristic. In some embodiments, it is possible to create an image of the biological object using both the first and second embodiments because the images provide complementary information.

[0033] According to one or more exemplary embodiments, the parameter represents a variance of the intensity of the interference signal over time.

[0034] According to one or more exemplary embodiments, calculating at least one pixel value for each elementary detector of the two-dimensional acquisition device includes calculating a Fourier transform of the interference signal acquired by the elementary detector as a function of time.

[0035] The present description relates according to a second aspect to a three-dimensional imaging system for imaging at least one transparent biological object in a biological sample by full-field optical tomography, said imaging system comprising: - a light source configured for emitting an illumination beam of spatially non-coherent light of a given central length, said illumination beam being configured to illuminate said sample during transmission; - an optical imaging system including a microscope lens with a given optical axis and a given object focal plane, near which the sample is positioned during operation; - means for displacing the microscope lens relative to the sample along an axial direction parallel to the optical axis of the microscope lens; - a two-dimensional acquisition device including a detection plane, said detection plane being optically conjugated with the object focal plane of said microscope lens by said optical imaging system; a processing unit; and for each section of the plurality of sections of the biological subject: - the three-dimensional imaging system is configured for the acquisition of a plurality of two-dimensional interference signals resulting from optical interference between the illumination beam and a beam scattered by the slice using the two-dimensional acquisition device; the processing unit is configured to calculate at least one first image of the sliced ​​field from the plurality of two-dimensional interference signals.

[0036] Such novel imaging system arrangements enable optical tomographic imaging of transparent biological objects, thereby making it possible to produce images of sections or "slices" of the biological object to create a three-dimensional image.

[0037] According to one or more exemplary embodiments, the two-dimensional interference signals of the plurality of two-dimensional interference signals are obtained for different positions of the subject focal plane within the thickness of the slice, resulting in a plurality of predetermined phase shifts between the illumination beam and the scattered beam ranging between -π / 2 and π / 2.

[0038] In this first embodiment of the imaging system according to the present specification, the means for displacing the microscope lens relative to the sample is configured not only to displace the object focal plane of the microscope lens within a slice of the biological subject to create an image of the slice, but also to displace the object focal plane within the sample along the optical axis to create a three-dimensional image to create images of multiple slices of the biological subject.

[0039] According to one or more exemplary embodiments, the calculation of the at least one first image comprises a linear combination of the two-dimensional interference signals of the plurality of two-dimensional interference signals.

[0040] According to one or more exemplary embodiments, the two-dimensional interference signal of the plurality of two-dimensional interference signals is acquired for a fixed position of the microscope lens relative to the sample, and calculating the at least one first image of the section comprises: - calculating, for each elementary detector of said two-dimensional acquisition device, at least one pixel value as a function of the value of a parameter representative of the temporal variation in the intensity of said two-dimensional interference signal acquired by said elementary detector.

[0041] In this second embodiment of the imaging system according to the present specification, the means for displacing the microscope lens relative to the sample can be configured only to displace the object focal plane along the optical axis within the biological object to be imaged in order to create images of multiple slices of the biological object such that a three-dimensional image is created.

[0042] The first and second embodiments may be implemented within the same imaging system according to the present specification to obtain a plurality of first images and a plurality of second images of a section of the transparent biological subject that is desired to be imaged.

[0043] Further advantages and features of the imaging techniques presented above will become apparent from the following detailed description taken in conjunction with the drawings identified below. [Brief description of the drawings]

[0044] [Figure 1A] 1 is a schematic diagram illustrating an example three-dimensional imaging system according to the present disclosure. [Figure 1B] 2 is a schematic diagram showing in greater detail the object focal plane of a microscope lens in an exemplary three-dimensional imaging system according to the present specification. [Figure 2A] A schematic diagram illustrating two positions within a slice of a transparent biological subject at the object focal plane of a microscope lens for creating an image of the field of the slice in an exemplary implementation of a three-dimensional imaging method according to one embodiment. [Figure 2B]FIG. 1 is a schematic diagram illustrating multiple positions of the subject focal plane of a microscope lens in an exemplary implementation of a three-dimensional imaging method according to a first embodiment, where the subject focal plane is centered on different slices of a transparent biological subject to create a three-dimensional image of the subject. [Figure 3A] 1 is a graph illustrating the theoretical intensity calculated at the center of an image field as a function of the relative position of the subject focal plane of a microscope lens to the position of a two-dimensional interference signal resulting from the interference of an illumination beam and a beam scattered by a scattering particle, such as in the case of a nanometer-sized intracellular organelle. [Figure 3B] 3B is a graph illustrating the theoretical intensity calculated and represented in FIG. 3A minus the average value of the interference signal in z corresponding to the continuous background. [Figure 3C] 1 is a graph showing, by way of example, the difference between the theoretical intensities calculated at the center of the field for each of two two-dimensional interference signals resulting from the interference of an illumination beam and a beam scattered by a particle for two positions of the particle separated by a distance equal to one-tenth of a given depth of field in an exemplary implementation of a three-dimensional imaging method according to a first embodiment, as a function of the relative position of the object focal plane of the microscope lens with respect to the position of the particle. [Figure 3D] FIG. 3C is a graph illustrating the theoretical intensity calculated and represented in FIG. 3B overlaid with a curve illustrating the relative displacement of the object focal plane of the microscope lens relative to the sample according to a sine function. [Figure 4] A schematic diagram illustrating the position of the object focal plane of a microscope lens relative to a slice of a biological subject for creating an image of the field of the slice in an implementation of a three-dimensional imaging method according to a second embodiment. [Diagram 5] FIG. 3C is a graph illustrating the theoretical intensities calculated and represented in FIG. 3B superimposed with a curve illustrating the random displacement of the essential structures of a biological object relative to the object focal plane of a microscope lens. [Figure 6]FIG. 1 shows experimental images obtained on two adjacent HeLa cells along two different section planes at a distance of 4 microns using two embodiments of the method according to the present specification. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0045] 1A and 1B generally illustrate an exemplary embodiment of a three-dimensional imaging system 100 according to the present disclosure.

[0046] The imaging system 100 includes a light source 110 configured to illuminate the sample 10 during transmission with an illumination beam of spatially incoherent light of a given central length λ. Advantageously, but not necessarily, the illumination beam is a temporally incoherent light beam to avoid undesirable optical speckle effects.

[0047] The light source 110 may be, for example, a light emitting diode (LED), a heated filament light source, a matrix of VCSEL (Vertical Cavity Surface Emitting Laser) type light emitting elements, a membrane associated laser, or any other means for making an illumination beam spatially incoherent.

[0048] In the example illustrated in FIGS. 1A and 1B, a light source 110 is positioned opposite the sample 10 to emit a transmitted illumination beam, although other illumination configurations are possible.

[0049] The three-dimensional imaging system 100 further includes an optical imaging system 120 that includes a microscope lens 121 with an optical axis Δ and an object focal plane 125 (FIG. 1B) near which the sample 10 is placed during operation. Additionally, in the example of FIG. 1A, the optical imaging system 120 includes an (optional) lens 122, commonly referred to as a "tube lens."

[0050] The three-dimensional imaging system 100 also comprises means for displacing the microscope lens 121 relative to the sample 10 along an axial direction parallel to the optical axis of the microscope lens. In the example of Figures 1A and 1B, the relative displacement means comprises a first piezoelectric element 131 configured to axially displace the microscope lens 121 and a second piezoelectric element 132 configured to axially displace a sample holder (not shown) on which the sample 10 is placed. The relative displacement means also comprises a control unit 135 for controlling the elements 131, 132. However, in other embodiments the relative displacement means may comprise only one element for displacing either the microscope lens or the sample holder.

[0051] The three-dimensional imaging system 100 further comprises a two-dimensional acquisition device 140 comprising a detection plane 141, i.e. optically conjugated by the optical imaging system 120 with the object focal plane 125 of the microscope lens, and a control module 145 of the acquisition device 140. The two-dimensional acquisition device 140 comprises a plurality of elementary detectors, or "pixels", arranged on the detection surface in the form of a matrix arrangement, for example in the form of a two-dimensional matrix arrangement. For example, the dimensions of the matrix arrangement of the elementary detectors define the dimensions of the image field 142 of the imaging system according to the present description. In another example embodiment, the dimensions of the image field 142 are limited by a field stop so that the effective detection surface or "ROI" (region of interest) of the detection plane is smaller than the surface covered by the two-dimensional array of the elementary detectors. The image field is rectangular, for example with dimensions between about 5 mm and about 50 mm.

[0052] For example, the 2D acquisition device is a CCD or CMOS type camera. Of course, other cameras can also be used, for example ultrafast cameras with a larger load capacity per pixel, for example cameras of the QUARTZ series by ADIMEC.

[0053] The control module 145 controls the acquisition by the acquisition device 140 , receives the electrical signals transmitted by the acquisition device 140 and sends them to the processing unit 150 .

[0054] As illustrated in FIG. 1A, the three-dimensional imaging system 100 includes a processing unit 150 configured for implementing the calculations and / or processing steps implemented in the methods according to the present application.

[0055] Generally speaking, when the specific implementation of a method step is referred to herein as a computational or processing step, it is to be understood that each computational or processing step may be implemented by software, hardware, firmware, microcode, or any suitable combination of these technologies. When software is used, each computational or processing step may be implemented by computer program instructions or software code. These instructions may be stored in or transmitted to a computer-readable storage medium (or processing unit) and / or executed by a computer (or processing unit) to implement these computational or processing steps.

[0056] Thus, in operation, as will be described in more detail below, the three-dimensional imaging system 100 is configured for acquisition, using said two-dimensional acquisition device 140, of a plurality of two-dimensional interference signals resulting from optical interference between an illumination beam incident on the sample 10 and a beam scattered by said at least one transparent biological object, e.g. a cell or a cluster of cells, of the sample 10. Additionally, the processing unit is configured to calculate, from the plurality of two-dimensional interference signals, one or more images of a field of view of each of the plurality of slices 101 of the biological object.

[0057] The lateral dimensions of the field are equal to the lateral dimensions of the image field divided by the magnification of the optical imaging system 120, which includes the microscope lens. Thus, for example, for a substantially rectangular image field with dimensions of 10 mm by 10 mm, the field dimensions are 100 μm by 100 μm if the optical imaging system has a magnification of 100×, and 500 μm by 500 μm if the optical imaging system has a magnification of 20×. Generally speaking, the field dimensions are between about 50 μm and about 500 μm.

[0058] The field of the slice 101 of the biological subject for which an image is to be formed can be defined by a volume consisting of a set of volume elements or "voxels", each voxel corresponding to a cylindrical volume with a length equal to the depth of field of the microscope lens and with a slice component defined by the diffraction disk of said microscope lens.

[0059] The depth of field L and the diameter φ of the intercept component are given by:

[0060]

number

[0061]

number

[0062] In this, n is the refractive index of the medium immersing the object space (e.g., in the case of an oil immersion lens, a medium with refractive index n≈1.5), NA is the numerical aperture of the microscope lens in said medium, and λ is the central wavelength of the illumination beam.

[0063] As shown in Fig. 1B, each slice 101 is substantially perpendicular to the optical axis Δ of the microscope lens. In operation, to create an image of the slice, the object focal plane of the microscope lens is centered on said slice using relative displacement of the microscope lens and the sample 10. From the images of the slices, a three-dimensional image of the transparent biological object can be created.

[0064] A first embodiment of a three-dimensional imaging method is explained with the aid of figures 2A, 2B, 3A, 3B, 3C and 3D. This first embodiment describes a first mode for acquiring and calculating images of a sliced ​​scene.

[0065] In an acquisition mode according to the first embodiment, the object focal plane of the microscope lens is displaced within a section of a transparent biological object, such as a cell or a cluster of cells, along an axial direction parallel to the optical axis of the microscope lens to create an image of the field of the section.

[0066] Multiple 2D interference signals are acquired for different object focal plane positions within the slice, resulting in multiple phase shifts between the incident beam and the beam scattered by the slice ranging between -π / 2 and π / 2. An image of the entire field of the slice, and thus the concept of full-field imaging, is then calculated directly from the 2D interference signals.

[0067] By way of illustration, FIG. 2A represents a cluster 20 of two cells 21, 22 for which it is desired to create a three-dimensional tomographic image. Each cell contains a nucleus 211, 221 with a nucleolus 212, 222 and a cytoplasm 213, 223 around the nucleus, which contains organelles 214, 224 such as mitochondria, vesicles, fat bodies, protein condensates, etc. The cytoplasmic organelles are located in the nucleus. The internal structure of the cytoplasmic organelles or the nucleus of the cell scatters light. As described below, it is possible to calculate an image of a slice in the volume of the cell from at least two two-dimensional interference signals obtained by varying the distance between the microscope lens and the sample, typically between about 1 / 10 and about 1 / 2 the depth of field, and thus varying the phase of the interference signals, for example by pixel-by-pixel difference between the two obtained interference signals. Dashed lines 201, 202 illustrate, by way of example, two positions of the object focal plane of the lens within the slice that is sought to be imaged.

[0068] To explain the calculation principle of the image of the section according to the first embodiment, we assume that the electromagnetic field received in the detection plane 141 of the acquisition device 140 (Fig. 1A) is a superposition of the incident field emitted by the light source and transmitted by the sample, modeled in this example by a plane wave, and the field scattered by the cells, modeled in each voxel of the field by a Gaussian beam suitable for analytical formulation. The conclusions developed below are also valid in the case of illumination beams that are not strictly plane waves or scattered beams that do not correspond strictly to Gaussian beams.

[0069] It should be noted that due to the spatial incoherence of the illumination beam, the beam scattered by the entire field can actually be thought of as multiple beams arranged side-by-side throughout the field of the biological subject, with the size of each beam being defined by the resolution of the microscope lens.

[0070] More precisely, the complex electric field of a Gaussian beam propagating along an axis z (Figure 1B) parallel to the optical axis of the microscope lens can be written in a scalar approximation as:

[0071]

number

[0072] The coordinate z is defined along the optical axis of the microscope lens relative to the origin (z=0) defined in the object focal plane of the microscope lens; r is the distance from the optical axis and k=2π / λ, λ being the central wavelength of the illumination beam. E0 and w0 are the field and size of the beam at the origin (z=0, r=0). It is possible to approximate the size of the beam at the origin w0 as being equal to the radius φ / 2 of the diffraction disk of the microscope lens given above [Equation 2]. It should be noted that in practice, the size of the beam defines the dimensions of the smallest accessible detail in the image. In order to obtain a sampling in the interference signal acquired by the detection device according to the sampling theorem, also known as Shannon's theorem, a value can be chosen for w0 that is greater than or equal to the maximum size of a pixel.

[0073] In what follows we give a proof around r=0, but in practice extended fields are involved and it is possible to show that the curves described below will be the same at any point in the field when the microscope lens is free of the relevant aberrations in practical use.

[0074] For a Gaussian beam, the width w(z) of the beam is minimal at its origin (z=0) and at a distance z along the beam axis, w(z)=w0(1+(z / z0) 2 ) 0.5 is equal to.

[0075] Depth of field is 2×z0=πw0 2 / λ et, and the radius of curvature of the wave is R(z)=z(1+(z / z0) 2 ).

[0076] From the wave electric field equation, it is possible to introduce a phase term, called the Gouy phase, which represents the phase jump occurring at the focus and is given by:

[0077]

number

[0078] Biological objects such as cells consist of internal structures with dimensions from about 10 nanometers to about 1 micrometer. For example, internal structures are intracellular organelles, filaments and microtubules, vesicles, and mitochondria. To illustrate the axial response of the recorded signals, we take here an example of such an internal structure that corresponds to a particle or "scatterer" having a given position (z) relative to the object focal plane of the microscope lens.

[0079] The field scattered by a particle in the center of the field is therefore described as:

[0080]

number

[0081] σ / A is the ratio between the effective scattering cross section of the particle and A, the area of ​​the incident illumination beam.

[0082] The incident field is modeled by a plane wave:

[0083]

number

[0084] Therefore, the intensity corresponding to the two-dimensional interference signal at the center of the field measured by the acquisition device is common to both beams and, neglecting the propagation and exp(iωt) terms which average to zero, is:

[0085]

number

[0086] That is:

number

[0087] Generally speaking, for small particles, the effective cross section is very small compared to 1 and can therefore be described as:

[0088]

number

[0089] Thus, FIG. 3A illustrates the calculated intensity curve I(z) for σ / A=0.01 and z0=0.25 μm (λ=0.5 μm, NA=1.25), i.e. a depth of field of 0.5 μm.

[0090] This property of the variation in the interference signal measured in the detection plane as a function of the position of the scattering particle is exploited to calculate an image of the field of a transparent biological object, for example a section of a cell.

[0091] As mentioned before, the position of the microscope lens relative to the sample is changed, for example, using a piezoelectric element (132 in FIG. 1A), and thus it is possible to vary the value of z and obtain several different interference signals.

[0092] More precisely, the linearity of the interference term in the vicinity of z=0 is exploited to obtain a useful cross-section of the object placed in the focus of the microscope lens (Fig. 3B). In this example, Fig. 3B represents the curve obtained by subtracting the average background from the curve shown in Fig. 3A.

[0093] Near the focus, the sine and arctangent functions are linearized to 1 / (1+(z / z0) 2 It is possible to remove the prefactor in (Eq. 9) so that the formula becomes:

[0094]

number

[0095] It can be seen that the linearity zone extends over a zone of the order of one micron on the curve illustrated in FIG. 3B.

[0096] Successive images for two positions with a z-distance of 0.1 μm are taken, e.g. by moving the lens (or the sample), and the difference between the two interference signals is measured, yielding:

[0097]

number

[0098] In addition, when the particle is far from the focus, sin(atan(z)) tends to zero as z gets large, and the prefactor (1 / (1+z 2 ) also tends to zero, so the strength of the interference term tends to zero.

[0099] As a result, when the particle is far from the focus, the difference between successive images shifted by Δz ∼ 1 μm becomes ΔI(z>>z0) = 0.

[0100] Therefore, the intensity difference will be non-zero only if the object is close to the focal point of the lens.

[0101] The imaging method according to the present specification therefore allows "optical sectioning", i.e. imaging of a slice of a sample of a given thickness. In other words, if a 3D object consisting of several scatterers / particles is looked at by image difference at two positions, only the scatterers present at a given depth are isolated and it is possible to obtain the structure of the 3D object at a given depth (its composition within the scatterers).

[0102] The phase shifts of -π / 2 and π / 2 occur at distances of the order of the depth of field z0 of the microscope lens. To obtain thin sections of a sample, lenses with large numerical apertures need to be used. For example, for an oil immersion lens with a numerical aperture of 1.25 the depth of field is 0.5 μm, and for water immersion lenses, often with a numerical aperture of 0.9 the depth of field is 1 μm.

[0103] In the general case of the above curves, it is possible to plot, pixel by pixel, the difference in intensity between two images taken at a distance of the order of the depth of field, e.g. 1 / 10 of the depth of field, for various positions of the scatterer, as illustrated in Figure 3C.

[0104] From the curve illustrated in Figure 3C, it is observed that the scatterer is only visible if it is placed close to the focus. It is possible to define an arbitrary criterion (for example full width at half maximum) to define the optical slice, i.e. the slice over whose thickness the scatterer is detectable, and here it turns out to be of the order of the depth of field of the device, i.e. of the order of 0.5 microns for an oil immersion lens with a numerical aperture of 1.25.

[0105] In practice, it has been decided that it is interesting to select interference signals obtained for axial differences between 0.1 and 0.4 microns.

[0106] In practice, for high rate operations, the piezoelectric element is not always adapted to provide an instantaneous square wave displacement, but it may be preferable to perform a sinusoidal modulation of the displacement. When in the zone where the optical response is linear over z~0, the fluctuations in the optical signal are also sinusoidal and can be very well filtered by synchronous detection employing, for example, four images.

[0107] If the scatterers fall outside this linear zone, it will not be possible to detect the signal by synchronous detection.

[0108] From the image of the optical slice, it is then possible to scan the sample within the volume, as illustrated in Figure 2B, followed by "tomography", i.e., measurements of the intensity scattered by the slice at various corresponding axial positions that provide a 3D image of the sample.

[0109] For example, the object focal plane can be positioned at 201 in Fig. 2B and interference signals can be acquired at this plane as well as at a plane 202 that is 0.2 micrometers away from plane 201. The difference between these two-dimensional interference signals is then calculated and an image of the first slice is determined, then the object focal plane is displaced and a new acquisition is made to determine images of the subsequent slices. For example, to calculate an image of another slice of light, the object focal plane can be displaced to plane 203, an interference signal is acquired in this plane and the difference with the interference signal acquired in plane 202 can be calculated, and so on.

[0110] The relative position of the microscope lens and the sample can also be periodically modulated by applying a sinusoidal voltage to one of the two piezoelectric displacements in Figure 1. Synchronous detection of the signal is then performed by time sampling of at least two images over half the period.

[0111] A second embodiment of the three-dimensional imaging method is described with the aid of figures 4 and 5. In this second embodiment, another mode for the acquisition and calculation of the images of the slices is described, which differs from the description in relation to the preceding figures. It should be noted that for each slice, it is possible to calculate a first image corresponding to the first embodiment and a second image corresponding to the second embodiment.

[0112] In the previous embodiment, a relative displacement of the microscope lens and the sample was performed and the difference in intensity at two different positions was measured.

[0113] However, applicants have found that most organelles in living subjects have internal movements that spontaneously modulate their positions.

[0114] Thus, in a second embodiment of the tomographic imaging method according to the present invention, as will be described in more detail below, there is no relative displacement of the microscope lens with respect to the sample for the calculation of the images of the light slices, which is only performed for the calculation of the images of the different light slices and thus for the reconstruction of the three-dimensional image.

[0115] However, in practice, since the displacements of biological subjects are inherently random, this displacement is not known in advance and is different from ΔI~σ.Δz [Equation 11], and therefore neither the effective cross section of the scatterer nor its displacement can be determined and it is not possible to characterize a single scatterer.

[0116] In the tomographic imaging method according to the second embodiment, an image of the field of an optical slice is calculated from the temporal variations in intensity between two-dimensional interference signals acquired for a given position of the object focal plane within the slice.

[0117] For example, for each elementary detector of a two-dimensional acquisition device, it is possible to calculate a pixel value as a function of the value of a parameter representative of the temporal variation in the intensity of the two-dimensional interference signal acquired by said elementary detector, the parameter representing for example the temporal variance of the intensity of the interference signal.

[0118] Indeed, the temporal fluctuations or variations in the intensity signal give information about the fluctuations in the position of the scatterers. For example, if the scatterers have Brownian motion (random fluctuations in position), then the measured intensity signal must be purely random, with ΔI that must statistically increase according to √t (or according to the square root of the time difference between the images). In practice, scatterers inside biological cells are more frequently considered. The physiology of the cells indicates that the cells control the movement of the scatterers (for example by molecular motors), so that the movement is no longer random but biased. By measuring the fluctuations in the intensity signal produced by the movement of the scatterers, it is possible to show that an intrinsic signal is accessible that depends on the metabolism of the cells.

[0119] In practice, N images as a function of time in the same plane are recorded and the signal is inherently random since the positions of the scatterers are also random, as illustrated in Figure 5.

[0120] From this time trace, statistical processing can be performed to provide information about the scatterer environment. The sensitivity to scatterer motion is only valid for scatterers within the optical slice, in other words very close to the focal plane. Signal fluctuations are therefore mainly caused only by scatterers coming from the focal volume.

[0121] It is thus finally possible to perform dynamic tomography, i.e. recording the signal fluctuations in one plane, analyzing them, moving the lens position and starting again - the scatterer environment is thus mapped in 3D.

[0122] For example, each pixel of the camera records a signal according to a time series that needs to be analyzed and represented. This analysis can be performed in several ways. For example, it is possible to calculate the Fourier transform of the time series of the interference signal and deduce from them a number of pixel values, for example values ​​corresponding to the power (integral over the spectrum of the squared modulus of FT), the center frequency (H) of the frequency spectrum, and the spectral width (S), respectively. For example, it is possible to use the HSV (hue, saturation, value) color representation to represent the center frequency (H), the spectral width (S), and the power (V), respectively, of the frequency spectrum.

[0123] Other means of analysis can be used, such as a simple calculation of the standard deviation of the interference signal time series, or an analysis of the cumsum function of said time series for biased behavior.

[0124] FIG. 6 represents an image of two sections of HeLa cancer cells separated by a distance of 4 microns obtained using an imaging system similar to that depicted in FIG. 1A according to the present specification. The microscope lens is an oil immersion lens of magnification 100× with a numerical aperture of 1.25 and a central wavelength equal to 450 nanometers. The camera used is a Photon Focus® MV-D1024E series CMOS camera.

[0125] To create this image, the object focal plane of the microscope lens was displaced and centered on two sections of the cell, the first one in the cytoplasm (61, 63) and the second one in the nucleus of the cell (62, 64). These images illustrate the ability of the method object here to tomograph cells. The lower images 63, 64 represent the morphology of the cell, the upper images 61, 62 the "dynamic" part. The scale bar equals 2 microns.

[0126] While described through a specific number of embodiments by way of example, it will be understood that the tomographic imaging methods and systems according to the present specification embrace various alternatives, modifications, and improvements that will be apparent to those skilled in the art, and which form part of the scope of the present invention as defined by the following claims. [Explanation of symbols]

[0127] 10 Samples 20 Clusters 21 cells 22 cells Images 61 and 62 63, 64 Images 100 3D imaging system, imaging system 101 Intercept 110 Light source 120 Optical Imaging System 121 Microscope Lens 122 Lens 125 Object focal plane 131 First Piezoelectric Element 132 Second Piezoelectric Element 135 Control Unit 140 2D acquisition device, acquisition device 141 Detection plane 142 Image Field 145 Control Module 150 Processing Units 201, 202 dashed line 211, 221 cell nucleus 212, 222 nucleolus 213, 223 cytoplasm 214, 224 intracellular organelles Δ optical axis λ central wavelength of the illumination beam

Claims

1. A method for three-dimensional imaging of transparent biological objects in a biological sample (10) by full-field optical tomography, comprising: positioning said sample in the vicinity of an object focal plane (125) of a microscope lens (121) encompassing a given optical axis (Δ); - illuminating said sample during transmission with an illumination beam of spatially incoherent light with a given central wavelength (λ); - displacing the microscope lens relative to the sample along an axial direction parallel to the optical axis of the microscope lens to define a number of positions of the sample, each position corresponding to a slice (101) of the biological object centered on the object focal plane of the microscope lens; - generating, for each position of said sample, at least one first image of the field of said section; and generating the at least one first image comprises: acquiring, by means of a two-dimensional acquisition device (140) comprising a number of elementary detectors arranged in a detection plane (141) resulting from optical interference between the illumination beam incident on the field and a beam scattered by the field, said detection plane being optically conjugated with the object focal plane of said microscope lens by means of an imaging optical system comprising said microscope lens; - calculating said at least one first image from said plurality of two-dimensional interference signals using a processing unit (150); A three-dimensional imaging method comprising:

2. the two-dimensional interference signals of the plurality of two-dimensional interference signals are acquired for different positions of the object focal plane within the thickness of the slice, resulting in a plurality of predetermined phase shifts between the illumination beam and the scattered beam ranging between -π / 2 and π / 2; The imaging method of claim 1.

3. said calculating said at least one first image comprises a linear combination of said plurality of two-dimensional interference signals. The imaging method of claim 2.

4. the relative displacement of the microscope lens with respect to the sample follows a periodic function with a maximum amplitude of λ / 4, where λ is the central wavelength of the illumination beam. The imaging method according to claim 2 or 3.

5. the two-dimensional interference signal of the plurality of two-dimensional interference signals is acquired for a fixed position of the microscope lens relative to the sample, calculating the at least one first image of the scene of the slice comprises calculating, for each elementary detector of the two-dimensional acquisition device, at least one pixel value as a function of the value of a parameter representative of the temporal variation in the intensity of the two-dimensional interference signal acquired by said elementary detector, The imaging method according to claim 1 or 2.

6. the parameter represents a temporal variance of the intensity of the interference signal; 6. The imaging method of claim 5.

7. A three-dimensional imaging system (100) for imaging transparent biological objects in a biological sample (10) by full-field optical tomography, comprising: a light source (110) configured for emission of an illumination beam of spatially incoherent light with a given central length, configured to illuminate said sample during transmission; an optical imaging system (120) comprising a microscope lens (121) with a given optical axis (Δ) and a given object focal plane (125), near which, in operation, said sample (10) is positioned; - means (131, 132, 135) for displacing said microscope lens relative to said sample along an axial direction parallel to said optical axis of said microscope lens; a two-dimensional acquisition device (140) comprising a plurality of elementary detectors arranged in a detection plane (141), said detection plane being optically conjugated by said optical imaging system with said object focal plane of said microscope lens; a processing unit (150), For each of the plurality of slices of the biological subject, - said three-dimensional imaging system is configured for said acquisition, by means of said two-dimensional acquisition device (140), of a plurality of two-dimensional interference signals resulting from optical interference between said illumination beam and a beam scattered by the field of said section, said processing unit (150) is adapted to calculate at least one first image of said scene of said slice from said plurality of two-dimensional interference signals; A three-dimensional imaging system (100).

8. the two-dimensional interference signals of the plurality of two-dimensional interference signals are acquired for different positions of the object focal plane within the thickness of the slice, resulting in a plurality of predetermined phase shifts between the illumination beam and the scattered beam ranging between -π / 2 and π / 2; 8. The imaging system of claim 7.

9. and wherein the calculation of the at least one first image comprises a linear combination of the two-dimensional interference signals of the plurality of two-dimensional interference signals.

9. The imaging system of claim 8.

10. the two-dimensional interference signal of the plurality of two-dimensional interference signals is acquired for a fixed position of the microscope lens relative to the sample, and calculating the at least one first image of the field of the section; - calculating, for each elementary detector of said two-dimensional acquisition device, at least one pixel value as a function of the value of a parameter representative of the temporal variation in the intensity of said two-dimensional interference signal acquired by said elementary detector; 9. The imaging system of claim 7 or claim 8, comprising:

11. the parameter represents the temporal variance of the intensity of the two-dimensional interference signal; The imaging system of claim 10.

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