Optical device, system and method for acquiring photoacoustic images using a homogenized beam from a pulsed light source - Patents.com

Abstract

The invention relates to an optical device (1) for shaping and homogenizing a light beam (2), comprising a light source (3) of a light beam or light pulses (2), a multimode optical fiber (7) having a core with a polygonal cross section, an optical coupling module (5) for guiding the light beam (2) from the light source (3) into the fiber (7), and a light direction module (13) for guiding and focusing the homogenized light beam or light pulses (14) from the outlet side (7") of the optical fiber (7) towards a sample (19). Also disclosed is a system using the device (1), further comprising an ultrasonic receiving transducer (21) and an acquisition and processing unit, for acquiring photoacoustic images of a sample (19) in optical or acoustic resolution photoacoustic microscopy and / or photoacoustic tomography applications.

Description

[Technical field]

[0001] The present invention belongs to the field of ultrasound imaging technology. More particularly, the present invention relates to devices, systems and methods for generating photoacoustic or optical acoustic images, which allow generating functional, molecular and anatomical images of substances or structures present in various tissues or samples, preferably of biological origin. [Background technology]

[0002] The photoacoustic effect describes the generation of ultrasound waves in tissues previously irradiated with short light pulses (understood as light beams) of high intensity, typically generated by a laser source. The molecules of the tissues on which these pulses (with durations in the nanosecond range and frequencies in the visible or near infrared bands) impinge absorb part of the energy of the light pulses and convert it into a series of rapid thermoelastic expansions, thus generating acoustic pressure waves in the ultrasound range.

[0003] The use of this effect, particularly in obtaining images of biological tissues, is well known in the art and has in recent years enabled the development of a variety of techniques for obtaining functional, molecular and anatomical images in a non-invasive manner that is not based on the use of ionizing radiation, which is particularly advantageous when studying living tissue.

[0004] Similarly, to observe a particular area of ​​the illuminated sample with the right contrast, it is necessary to tune the light frequency to tune to the absorption band of the chromophores present there (areas present in the molecules of a particular substance where the energy difference between the molecular orbitals is in the visible or infrared spectral range). At this frequency, these areas that can re-emit the radiation absorbed by ultrasound have a greater light absorption than the surrounding substances, thus presenting a higher level of contrast in the final image. The chromophores used are usually endogenous, i.e. naturally present in the tissue studied. Examples of this type of substance include hemoglobin in blood vessels, melanin responsible for the pigmentation of the epidermis, some lipids in adipose tissue, collagen and elastin in connective tissue, etc. These substances, being a natural part of the sample studied, make it possible to observe the functional reactions of tissue molecules without the use of external chemicals, in contrast to conventional ultrasound, which only provides anatomical images, or other techniques such as more advanced functional ultrasound imaging methods (referred to as fUS), which require external contrast agents (such as microbubbles) to perform functional studies of the sample and mainly by measuring blood flow using the Doppler effect.

[0005] Despite the advantages of endogenous chromophores, in some cases exogenous chromophores can also be used, which are externally introduced into the sample with the aim of improving the photoacoustic properties. Examples of these compounds are gold nanoparticles or photosensitive contrast agents based on covalent organic frameworks (COFs), which have the advantages of being wavelength tunable, having a high degree of biocompatibility, being able to bind to specific organic molecules and be consumed by the tissue of special interest, as is the case with glucose in the study of malignant tumors.

[0006] Regarding the optical beam used to excite tissues so that they can emit ultrasound waves, the most common state-of-the-art approach is the use of short laser pulses with sufficient energy per unit area (optical fluence) to allow the generation of broadband ultrasound waves with a peak pressure proportional to the optical fluence absorbed by biological tissue. Optimal laser pulses for use in photoacoustic imaging techniques have pulse durations of a few nanoseconds to a hundred nanoseconds and a power of 10-100 mJ / cm, taking into account that the fluence is limited to values ​​below the standardized maximum permissible exposure value to avoid damage to biological tissues. 2 It has a fluence in the range of

[0007] For the source of these pulses, the usual choice is the use of solid-state laser devices. These lasers, based on Q-switching technology such as Nd:YAG, have high pulse energies of 200-500 mJ or more at two fixed wavelengths, 1064 nm and 532 nm, but are limited in terms of pulse repetition rates that vary from about 10 to 100 Hz. Furthermore, in the first place, some tunable lasers, such as optical parametric oscillators based on Nd:YAG, are usually used to access multiple wavelength excitations in the visible-near infrared band up to 2500 nm, but with the disadvantage of being performed with a rather low energy per pulse. However, due to the relatively high energy of the pulses generated by solid-state lasers, they are often preferred for photoacoustic tomography, since they can cover a larger excitation area or field of view with a fluence sufficient to activate detectable ultrasound.

[0008] On the other hand, photoacoustic microscopy (PAM) scanners attempt to obtain images with high spatial resolution by focusing a light beam on an extremely small excitation point and scanning the beam pixel-by-pixel across the entire viewing area and at various focal depths within the investigated volume. In particular, optically resolved photoacoustic techniques require much less pulse energy (on the order of microjoules or less) to generate photoacoustic waves detectable by common ultrasound transducers, since the light energy is concentrated in a small spot area. Therefore, low-energy light sources such as pulsed laser diodes (PLDs) become a feasible alternative for these imaging methods. Compared to solid-state lasers, PLDs offer a significantly higher pulse repetition rate of over 10 kHz, thus accelerating photoacoustic imaging, and are significantly cheaper and more compact, so they can be used to build portable and cost-competitive instruments.

[0009] On the other hand, the use of PLD in this field is accompanied by some drawbacks. In particular, semiconductor-based laser cavity emitters generate a highly divergent beam, so that the light beam obtained with this type of source is of low quality. Moreover, the energy generated by these sources is generally low, so that a common method uses multiple stack emitters, resulting in an extremely wide emission area and a non-uniform intensity pattern, so that the final beam is composed of many transverse modes. The multimode nature of the beam thus makes it extremely difficult to focus it on the point of interest, which is generally solved by using focusing lenses of high numerical amplitude. Although these lenses help to focus the beam on the relevant area, they also greatly reduce the working distance (WD). This means that the sound waves generated in the sample by the photoacoustic effect can only be generated at extremely short focal lengths by the optical devices usually used, thus limiting the depth at which photoacoustic images can be obtained in the tissue. This is one of the fundamental advantages of this technique compared to the rather superficial purely optical imaging techniques.

[0010] On the other hand, in laser diode-based photoacoustic microscopy devices, the multimode configuration of the beam does not allow for a reduction in the lateral spatial resolution without energy loss on the point of interest, resulting in a degradation of the quality of the final image, resulting in a poor signal-to-noise ratio, or in other words, a reduction in the contrast obtained at the ultrasound detector.

[0011] Some of these drawbacks are described in a general way in US Patent Application No. 2009109698A1. It describes a method for homogenizing or uniformizing a beam with an irregular transverse radiation profile, typically emitted by a laser diode, using a multimode optical fiber with a polygonal cross section, and also presents various methods for focusing the uniformized beam in a spatial light modulator for a projection or "display" device. Similarly, Japanese Patent Application No. 2003121664A, European Patent Application No. 2228674A1, and US Patent Application No. 2013170806A1 use the same concept of transmission through a square, rectangular, or other core fiber to solve the homogenization of beam fluence in other applications, for example, marking, photocoagulation, semiconductor lithography, etc. However, none of these documents are related to photoacoustic imaging applications.

[0012] In the field of photoacoustics, Chinese Patent Application No. 102854142A proposes the use of laser diodes as photoacoustic excitation sources using a simple beam focusing optical configuration without optical fiber and without homogenization of the laser diode beam. US Patent Application No. 2011303015A1 and Chinese Patent Application No. 101813672A propose beam homogenization methods based on the use of non-laser diode type laser sources with high quality Gaussian beam profiles and fundamental modes, but are significantly different because they do not use optical fibers. In the first case, a diffuser or holographic plate is used to uniformly expand the illuminated area, but not in a focusing manner. In the second case, beam shaping is used for large area diffuse illumination of the object and detection by a depth scanning piezoelectric detector array.

[0013] Finally, WO2009055705A2 describes a confocal photoacoustic microscope device (i.e. focused light beam) with several possible configurations. The described method uses a single mode optical fiber as a possible configuration for transmitting a Gaussian fundamental mode beam generated by a solid-state laser source such as Nd:YAG. Furthermore, spatial filtering is performed using a diaphragm or pinhole with the aim of removing higher order modes that the main Gaussian beam may contain.

[0014] For a more comprehensive understanding of the state of the art on laser diode-based photoacoustic imaging, readers can refer to the following literature: [S. Jeon, J. Kim, D. Lee, JW Bak, C. Kim, “Review on practical photoacoustic microscopy”, Photoacoustics 15 (2019) 100141], [M. Erfanzadeh, Q. Zhu, “Photoacoustic imaging with low-cost sources; A review”, Photoacoustics 14 (2019) 1-11] and [Q. Yao, Y. Ding, G. Liu, L. Zeng, “Low-cost photoacoustics imaging systems based on laser diode and light emitting diode excitation”, Journal of Innovative Optical Health Sciences Vol. 10 No. 4 (2017) 17300038].

[0015] In view of the problems raised with regard to the standardization of the beam emitted from a laser diode in photoacoustic imaging applications, regardless of the working distance (WD) or the depth at which said image can be generated in the tissue, the present invention proposes an improved method for generating images, not present in the prior art, which solves the aforementioned problems and allows the acquisition of high-quality images in photoacoustic microscopy techniques. Summary of the Invention [Means for solving the problem]

[0016] In order to overcome the above-mentioned limitations of the state of the art, the present invention aims to provide a novel method for obtaining low-cost photoacoustic images, which solves the problem of the non-uniformity of the light beam, thereby making it possible to increase the lateral spatial resolution of the photoacoustic images by using low-quality beam sources (such as laser diodes or LEDs) while exciting the sample under study, without reducing the focal length of the light beam and without reducing the maximum depth at which ultrasonic sound waves are generated in the sample under study by said light beam through the photoacoustic effect.

[0017] More specifically, a first object of the present invention relates to an optical device for shaping, homogenizing, guiding and focusing a pulsed light beam onto a material sample, comprising: a source of light beam or light pulses; A multimode optical fiber having a polygonal cross-section core with an optical inlet side and an optical outlet side; an optical coupling module configured to guide a light beam or light pulses from the light source to an input side of the optical fiber; a light direction module configured to direct and focus the shaped and homogenized light beam or light pulse from the exit side of the optical fiber onto a sample.

[0018] After passing through the light coupling module and the optical fiber, the beam from the light source, already shaped and homogenized in this way, is guided to a focal point by means of a light direction module for optoacoustic excitation, where known collimation and focusing optics can be used. These optics optically couple the exit plane of the fiber with the focal or focusing plane of the beam and reproduce the shape of the fiber core or the shape of the aperture placed at the end of the fiber. The zoom factor, or magnification, of the optical configuration used ultimately determines the enlargement or reduction in the size of the laser excitation focal cross section of the optoacoustic device.

[0019] Furthermore, in a preferred embodiment of the present invention, the light source is a pulsed laser diode, a light emitting diode (LED), or a stacked array thereof. The present invention enables the use of said light sources to generate high resolution photoacoustic images, taking advantage of their advantages, such as light weight or low cost, and overcoming these limitations by shaping and homogenizing the beam through the use of the device of the present invention.

[0020] In another preferred embodiment of the invention, the device comprises a surface with an aperture arranged at the exit side of the optical fiber, said aperture being arranged substantially centrally with respect to the core of said optical fiber. This allows spatial filtering of the light beam, resulting in a reduction in the size of the focal spot on the sample, which results in a significant improvement in the lateral resolution. Furthermore, the aperture, i.e. the aperture size, is mechanically adjustable, allowing the selection of the lateral resolution required for each application.

[0021] In another preferred embodiment of the present invention, the multimode optical fiber having a polygonal cross-section core has a high numerical aperture, which is particularly advantageous when a light source providing a highly divergent light beam is used, since a high numerical aperture allows the fiber to accept light beams coming from a wide range of incidence angles.

[0022] In another preferred embodiment of the invention, the light coupling module and / or the light direction module comprises two movable plano-convex aspheric lenses, one for collimation and the other for focusing. Furthermore, the direction module may comprise one or more motorized scanning mirrors, which allow precise focusing on a required point on the sample and then scanning the sample to obtain a complete image of it.

[0023] In another preferred embodiment of the invention, the optical device is part of a system for generating a photoacoustic image of a material sample. The system further comprises an ultrasonic receiving transducer configured to convert acoustic waves into electrical signals, and a signal acquisition and processing unit. Furthermore, the ultrasonic transducer is preferably ring-shaped and coaxially surrounds the optical axis of the light beam, and the signal acquisition and processing unit preferably comprises a preamplifier, an analog-to-digital converter, and a device with processing capabilities, such as a computer, which can be connected to a screen or display for real-time display of the photoacoustic image.

[0024] The presented system offers several technical advantages with respect to three major aspects of photoacoustic microscopy device performance, which are described below.

[0025] First, by pre-concentrating the energy of the laser beam in the core of the optical fiber, an increase in the laser fluence at the focal point is obtained, generating an acoustic pressure amplitude that is directly proportional to the fluence. In this context, the fluence of the laser beam homogenized by the optical fiber has an approximately constant cross-sectional profile in the illuminated area of ​​the focal plane. Thus, an acoustic resolution photoacoustic microscope (AR-PAM) device can provide uniform laser excitation over an illuminated area that bounds the field of view, with the lateral spatial resolution being determined by the focus of the ultrasonic transducer.

[0026] Second, in an optical resolution photoacoustic microscope (OR-PAM) setup, improved lateral spatial resolution of two-dimensional photoacoustic images is obtained, which is determined by the size of the cross-sectional area of ​​the laser focus: laser fluence (f) and focal area (A 0 ) is the larger or smaller polygon core area (A 1 ) can be tuned to the specific requirements of the photoacoustic microscope system by coupling more or less energy from the laser source light beam into the fiber entrance. The average fluence at the focus is then φ=E 0 / A 0 and E 0 A 0 =m d 2 ×A 1 is the energy of the light beam pulse that ultimately reaches a focus of area determined by A 1 is the cross-sectional area of ​​the fiber core, m d is the magnification factor provided by the focusing optics of the light direction module. For example, the increment m d For =1, the focal area corresponds to the cross-sectional area of ​​the fiber core.

[0027] Finally, the lateral spatial resolution improves regardless of the focal depth or associated working distance of the directional optics, in other words the focal area A 0 is reduced, which determines the maximum available displacement range of the laser focus to acquire two-dimensional images at various depths within the tissue, thus enabling three-dimensional or volumetric photoacoustic scanning imaging.

[0028] Due to the above mentioned technical advantages, the use of the devices and systems of the present invention for generating photoacoustic images is of particular interest in optical resolution photoacoustic microscopy applications, acoustic resolution photoacoustic microscopy applications, and / or photoacoustic tomography applications.

[0029] In addition to the introduced apparatus and system, which can be advantageously adapted to certain aspects of a low quality multimode pulsed light source beam applied to excite ultrasound in a photoacoustic imaging system, a method or procedure for generating a photoacoustic image of a material sample is also disclosed, which comprises performing the following steps by operating a system as described above: a) placing a light source (3) configured to emit a light beam or light pulse with a frequency in the visible or near infrared band, with a duration in the nanosecond range, at a distance from the sample that is equal to or shorter than the length of the optical fiber. b) guiding the light beam or light pulse coming from said light source to the entrance side of a high numerical aperture multimode optical fiber having a polygonal cross-section core by adjusting the positions of the elements constituting the optical coupling module. c) focusing said homogenized light beam or light pulse at a point on the sample by adjusting the positions of the elements constituting the light direction module. d) placing an ultrasonic receiving transducer on the surface of the sample (19); e) Electronically receiving and processing the signals captured by the transducer using a signal acquisition and processing unit to obtain an optoacoustic image.

[0030] Finally, steps c to e of the method are repeated at least once, changing the position of the focus of the focusing lens by a second adjustment of the position of the elements constituting the light direction module, generating pixels each time corresponding to different points of the sample in a plane perpendicular to the direction of the beam, and a subsequent step of assembling said pixels is performed according to their respective corresponding illumination positions of the sample to obtain a two-dimensional image of the sample. This same method for obtaining a two-dimensional image is repeated in sequence, so that in each repetition the position of the focusing lens relative to the ultrasonic transducer changes along the beam direction, and photoacoustic images of different planes of the sample are generated each time longitudinally in the beam direction, thus making it possible to obtain a three-dimensional image of the sample after performing a subsequent step of assembling said images according to their respective corresponding illumination positions of the sample.

[0031] The present invention describes a method for producing high quality photoacoustic microscopy images by using a pulsed laser diode or a stacked array thereof having a homogenized and shaped multimode beam that exhibits the following characteristics: Flat and uniform distribution of beam energy and fluence, providing high image contrast. High lateral spatial resolution down to a few micrometers. Working distance in the centimeter range, which allows access to a wide range of axial excitation depths within the sample and provides more space for placing the ultrasound detector.

[0032] In the present invention, these beam characteristics are achieved by performing two main steps. First, the light beam is coupled into a multimode optical fiber with a polygonal cross section (e.g. square, rectangular, octagonal, etc.) using a conventional optical coupling device known in the state of the art, for example, two movable plano-convex aspheric lenses (one for collimation, the other for focusing) and an optical fiber entrance support, or by direct coupling. The fluence and intensity pattern emitted by the PLD source is thus homogenized due to the mixing of the spatial modes of the original beam carried out in the fiber by multiple total internal reflections at the walls of its core. Furthermore, better beam fluence homogenization results are obtained when the fiber length is 1 m or more. This gives rise to a beam cross-sectional profile with a uniform and flat or flat-top intensity distribution at the exit side of the fiber. The cross section of the beam has the same shape as the fiber core, and the asymmetric divergence of the fast (vertical) and slow (horizontal) modes is symmetrized by the common numerical amplitude of the fiber. Similarly, lateral resolution is improved due to a diaphragm-type aperture or pinhole placed on the exit side of the optical fiber.

[0033] Other objects of the invention relate to the specific embodiments described in the claims of this application.

[0034] Within the meaning of the present invention, the term "substantially" should be understood to include "the same" or within a 10% variation range. [Brief description of the drawings]

[0035] These and other features and advantages will be more fully understood from the detailed description of the invention and preferred exemplary embodiments thereof, which are considered in conjunction with the accompanying drawings.

[0036] [Figure 1] It corresponds generally to a photoacoustic microscope optical resolution scanner, presented as a possible embodiment of the invention, based on the use of a PLD set as a laser light source coupled to a beam homogenizing device of the invention. [Figure 2a]1 shows a longitudinal view of a multimode optical fiber having a polygonal cross-section core for use in the method of the present invention; [Figure 2b] 1 shows cross-sectional views of two multimode optical fibers with polygonal cross-section cores (one square, one octagonal) for use in the method of the present invention. [Figure 2c] 1 shows the exit side of a multimode optical fiber having a polygonal cross-section core used in the method of the present invention, more specifically the location of a diaphragm opening or pinhole. [Figure 3a] An experimental CCD camera image with a pixel size of 4.5 micrometers and the beam cross-sectional fluence profile at the emission plane of a 3×3 PLD array with a total area of ​​200×280 square micrometers are shown. [Figure 3b] Experimental CCD camera images with a pixel size of 4.5 micrometers and the beam cross-sectional fluence profile in the emission region of the PLD focused at the entrance of a multimode fiber with a square fused silica core with 150 micrometer edges are shown. [Figure 3c] An experimental CCD camera image with a pixel size of 4.5 micrometers and the fluence profile of the homogenized beam cross section at the exit side of the optical fiber after passing through a 2 meter length of fiber are shown. [Figure 4a] Experimental CCD camera images with a pixel size of 4.5 micrometers and the beam cross-sectional fluence profile at the focus of the beam direction module for a square focus with 75 micrometer edges, homogenized with an almost flat profile are shown. [Figure 4b] The cross-sectional fluence profile of the beam at the focus of the optical direction module of the beam for a circular focus of 10 micrometers in diameter after spatial filtering using a diaphragm or pinhole at the exit side of the fiber is shown.

[0037] (See numbers on the drawing) (1) An optical device for shaping and homogenizing a beam (2) Light beam or light pulse (3) Light source (4) Controller (5) Optical coupling module (6) Focused pulse or beam (7) Optical fiber (7') Fiber inlet side (7”) Fiber Exit Side (8) Protective outer coating (9) Core coating (10) Fiber core (11) Apertures, diaphragms or pinholes (13) Light Direction Module (14) Shaped and homogenized light beam (15) Focus (16) Directional focusing lens (17) Motorized scanning mirror or mirror system (18) Image or imaging focal plane (19) Sample (20) Photoacoustic waves (21) Ultrasonic transducer (UST) (22) Preamplifier (23) Analog-to-digital converter (24) Processor or computer (25) Screen or display (26) Optical fiber entrance support (27) Combination collimator lens (28) Combining lens (29) Directional collimator lens DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0038] A detailed description of the invention is provided below based on Figures 1 to 4 of this document. This description is provided for illustrative purposes, without limiting the claimed invention.

[0039] 1 shows a schematic representation of the components of a system for generating an optoacoustic image, corresponding to an optoacoustic imaging scanner of the OR-PAM type, according to a preferred embodiment of the system of the invention. This scanner is divided into two main parts: an optical device (1) for the generation, shaping and homogenization of a light beam or light pulse (2) and an ultrasound detection and acquisition device. As well as the system for generating an optoacoustic image, the optical device (1) for the generation, shaping and homogenization of a light beam or light pulse (2) contained in said system is itself a main object of the invention.

[0040] In the optical device (1), a light beam or light pulse (2) is generated by a light source (3), preferably a PLD laser, via a current pulse from a controller (4). An optical coupling module (5) couples the power of the light beam (2) into an optical fiber (7) and converts said light beam (2) into a focused beam (6) at the entrance side (7') of the optical fiber. The optical fiber (7) is preferably of multimode type and of any length. Furthermore, the optical fiber (7) is preferably provided with an outer protective coating (8), a core coating (9) and a polygonal cross-section core (10). Figures 2a and 2b show the geometric details of a square and an octagonal optical fiber (7) as two examples for the proper shaping and homogenization of the PLD multimode light beam (2). Figure 2a shows a longitudinal cross-section of the optical fiber (7), showing its main components: the inner core (10), the outer protective coating (8) and the core coating (9). FIG. 2b also shows details of two examples of polygonal fiber cross-sections (7) with square core (10) and octagonal core (10) geometries.

[0041] Finally, the light direction module (13) guides the shaped and homogenized beam (14) from the fiber exit side (7") to a focal point (15) with an area size determined by the light direction module (13) where it is focused in an image plane (18) at a working distance of a focusing lens (16) which is directly related to the focal length. Scanning of the image plane (18) located at a specific penetration depth in the sample is performed using a system of motorized mirrors (17) installed in the light direction module (13) (shown in Figure 1), which moves the beam focal point (15) to different x-y coordinate positions of said image plane (18).

[0042] In addition to this, improved lateral resolution can be achieved by spatial filtering with a small aperture (11), e.g. a sheet or surface with a diaphragm or pinhole (12), centrally placed in front of the exit side (7") of the fiber (7) core, as shown in Figure 2c. In this figure, the exit side (7") of the fiber (7) is seen with the aperture (11) centrally placed. This is of particular interest for optical resolution photoacoustic microscopy setups. A nearly uniform fluence cross-sectional area at the exit of the fiber (7) allows the diaphragm (A' 1 The shape of the light beam (14) passing through an aperture area of ​​A' 0 ) into the focal plane (18) of the directional module (13) with a corresponding ratio A' 0 =m d 2 ×A' 1 According to the magnification factor m of the light direction module (13), d(values ​​less than or equal to 1), a better lateral resolution is obtained with an almost constant fluence given by the homogenized light beam (14) of the optical fiber. Using this spatial filtering method, the focus (15) and the lateral resolution can be reduced by decreasing the aperture of the diaphragm (11) to a few micrometers, depending on the laser wavelength used, without changing the effective focal length of the objective lens of the directional module (13), thus preserving the optical depth and working distance of the image. Furthermore, if the aperture of the diaphragm or pinhole (11) included in the optical device is mechanically adjustable, this offers an excellent advantage in selecting the desired lateral resolution for observing the sample.

[0043] Photoacoustic waves (20) are generated as ultrasonic waves in the region of the focal point (15) due to the photoacoustic effect as a result of the optical power (typically of a few nanoseconds duration) of the incident light beam (2), which is expressed by the well-known relationship P(r)=Γμ a The optical fluence at a given spatial point r in the sample (19) is described by φ(r), which describes the conversion of the optical fluence at a given spatial point r in the sample (19) into a pressure pulse of proportional amplitude P, where μ a is the optical absorption at this point and Γ is the Gruneisen parameter, which depends on the thermodynamic and elastic properties of the medium.

[0044] On the other hand, the ultrasound detection and acquisition device of the present invention preferably comprises an ultrasound transducer (UST) (21) coaxially surrounding the light beam (2) and arranged in a so-called reflection mode, either in direct physical contact with the sample (19) or through an intermediate highly ultrasound-transmitting medium (such as a coupling gel). The UST (21) receives the ultrasound waves (20) propagating in the medium at the speed of sound and converts them into electrical signals. These signals are amplified and digitized by a preamplifier (PREAMP) (22) and an analog-to-digital converter (ADC) (23), respectively, for subsequent digital processing of the image on a computer (24), as shown in FIG. 1. This computer can then be connected to a screen or display (25) for observing the generated images, sometimes in real time. Thus, in an OR-PAM scanner, each photoacoustic image frame is finally recorded as a 2D map of sound pressure amplitude (or squared amplitude), and each position of the laser focus (15) in the scanned image plane (18) corresponds to a pixel of the two-dimensional photoacoustic image.

[0045] Also, depth scanning can be performed along the Z-axis, for example by moving the focusing lens (16) relative to the UST (21) which is fixed and in contact with the surface of the sample (19), generating a 3D volumetric image of 2D frame slices. The maximum scanning range within the material sample (19) is provided by the working distance of the focusing lens (16), so that the focal spot (15) can be focused up to a maximum depth given by the working distance of the lens (16) from the sample surface. For this reason, a sufficient working distance is usually required to acquire images at depths up to a few millimeters, which is one of the main advantages of photoacoustic imaging compared to purely optical imaging techniques, which are much shallower. Furthermore, the use of a flexible optical fiber (7) allows an advantageous design, since the light direction module (13) can be moved independently and placed at any position on the sample (19) surface, facilitating the actual response for imaging.

[0046] Regarding the optical method of beam shaping in the OR-PAM scanner, a light beam or light pulse (2) is generated by a stack of PLDs used as light source (3) connected to a current pulse controller (4) or PLD driver via positive and negative terminals of DC polarity. The PLD controller (4) provides a train of pulses, which have a pulse width t programmed in the controller (4). p duration and pulse repetition rate F p Determined by D=t p F p When the PLD exceeds the current intensity threshold, it emits laser light with a duty cycle given by: The PLD then produces a pulsed light beam (2) with the same temporal profile and with the power of the light pulses proportional to the intensity of the current supplied to the PLD by the controller (4). To prevent damage to the PLD, the current intensity, duty cycle and pulse width are set below the maximum values ​​defined in the PLD manufacturer's specifications.

[0047] To achieve a higher maximum optical power, the PLD light source (3) consists of multiple individual laser diode emitters, typically arranged in horizontal monolithic bars with a certain number of emitters per bar, while these bars have a global emission area A 2 Thus, the cross-section of the emitted light beam (2) will have a non-uniform intensity or fluence pattern corresponding to any configuration of the array of individual emitters.

[0048] Figure 3a shows radiation area A 2 =200×280mm 2 Figure 3a shows an image (encircled by a black square in Figure 3a) taken by a CCD camera of the emission cross section of the light beam (2) of a particular stacked array of 3 × 3 PLDs with a 100 ns light beam width (2) and a 100 ns light pulse repetition frequency (RPF) of up to 10 kHz. For this PLD, the maximum power and central wavelength are about 200 W and 905 nm, respectively. The light beam (2) energy of this PLD beam is estimated to be 20 mJ with a 100 ns light beam width (2), and the repetition frequency of the light pulses (2) can reach up to 10 kHz.

[0049] As shown in the embodiment of the OR-PAM scanner in FIG. 1, a light beam or light pulse (2) from a PLD light source (3) directs most of the power of the light beam (2) into a PLD emission area A. 2 From the fiber core (10) region A 1 The optical fiber (7) is coupled to the optical fiber (7) having a polygonal cross-section core (10) using an optical coupling module (5) capable of launching the PLD radiation. In the case where the area of ​​the fiber core (10) is larger than the area of ​​the PLD radiation (A 1 >A 2 ), after proper alignment of the fiber core (10) with respect to the center of the PLD region, direct coupling can be implemented without optics. 1 =m c 2 ×A 2 Therefore, in order to match the cross section of the light beam (2) at least up to the area of ​​the fiber core (10), m c In order to achieve this, an optical coupling module (5) with a magnification factor of 0.015 μm may be required. Alignment of the position of the entrance side (7') of the fiber core (10) with the position of the beam (6) focused by the optical module on the fiber side (7') is usually required to maximize the entrance power delivered to the fiber core (10), which can be achieved, for example, by precisely adjusting the position of the fiber entrance support (26) with micrometer translations in two directions transverse to the optical axis.

[0050] The combined light beam (2) is guided along the fiber (7) until it emerges from the exit side (7") of the fiber (7) with a divergence given by the numerical aperture of the fiber (7) and an approximately uniform or homogenized fluence cross-section.

[0051] The numerical aperture (NA) of the fiber (7) provides the maximum acceptance angle of an incident light beam (2) that is guided through the fiber core (10) by multiple total internal reflections at the core / coating interface. The numerical aperture is defined by the well-known ratio NA=√(n core 2 -n coating2) based on the refractive index of the aforementioned fiber (7) elements. Thus, a highly divergent light beam (2) coming from the PLD light source (3) and the optical coupling module (5) typically requires a fiber (7) with a high numerical aperture to achieve effective mode mixing along the fiber core (10) and thus to obtain a better homogenization of the light beam (2) at the exit of the fiber (7). The length of the fiber (7) also improves the homogenization of the light beam (2), typically with a nearly constant fluence profile for a 1 or 2 meter fiber (7).

[0052] Figures 3b and 3c show CCD camera images of the cross-sectional fluence distribution of the light beam or light pulse (2) at the entrance of the fiber (7) in the focal plane of the light coupling module (5) and at the exit of the fiber (7), respectively. These experimental data are provided to evaluate the performance of the fiber homogenization (7), for the specific embodiment of the system of Figure 1, which is a combination of the previously described PLD stack with a 3x3 emitter configuration (see pattern in Figure 3a), a multimode optical fiber (7) with a square core (10) of 150 μm per side, a numerical aperture of 0.39, and a length of 2 m, and a light coupling module (5) consisting of two plano-convex aspheric lenses, the first lens (27) collimates the beam and the second lens (28) focuses it into the fiber core, with individual nominal focal lengths of f 1 =20.1mm, f 2 = 10.5 mm. Thus, this optical coupling module (5) corresponds to m c =f 2 / f 1 ≈0.5) and the PLD emission region A 2 is the cross section A of the fiber core (10) 1 The ray tracing simulations reproducing the same optical configuration show a good homogenization of the cross section of the fluence of the light beam (2) at the exit of the fiber (7) both for a square core (10) of 150 μm per side and for an octagonal core (10) with a virtual diameter of 150 μm (diameter of the octagon within the circumscribed circle).

[0053] The homogenized beam (14) emerging from the exit of the fiber (7) is brought to the final focus (15) by the light direction module (13) of the OR-PAM system or scanner. It maintains the collimation lens (29) and the focusing lens (16) and the xy scanning mirror system (17) on the same mechanical support, to which is connected a ring-shaped ultrasonic transducer (21) that coaxially surrounds the light beam (14), as shown in the embodiment of FIG. 1. In this optical configuration, the region A 1 The cross section of the beam (14) fluence formed at the exit of the fiber (7) has a region A 0 =m d 2 ×A 1 It appears with the same uniform fluence distribution throughout, the size of which depends on the ratio of the focal lengths of the collimation lens (29) and the focusing lens (16) and is determined by the magnification factor or rate of increase m d =f 2 / f 1 is proportional to.

[0054] FIG. 4a shows a CCD camera image of the cross-sectional shape of the laser focus (15) and the homogenized fluence distribution obtained for the same 150 μm square core (10) fiber (7) of the experimental example described above. The light beam (14) formed at the exit of the fiber with the fluence cross-section shown in FIG. 3c has a magnification factor or rate of increase m d =f 2 / f 1 The light is then brought to the focal plane (18) using a specific light directing module (13) consisting of two plano-convex aspheric lenses with focal lengths of f ≈ 0.5, where the focal lengths of the first lens (29) and the second lens (16) are f 1 =20.1mm, f 2= 10.5 mm. The final focus (15) thus has the same square shape as the fiber core (10), but with an area that is reduced by a factor of four with respect to the area of ​​the core (10) at the fiber exit, and the sides of the square area are reduced by a factor of two to about 75 μm per side. The maximum penetration range of the focus (15) into the sample is given by the working distance of the collecting lens (16), which is in particular the focal length f 2 = 10.5 mm, which gives a working distance in air WD = 6.4 mm and with the lens (16) immersed in water WD = 8.8 mm.

[0055] Since the size of the laser focus (15) directly determines the lateral resolution of the photoacoustic image formed by scanning the laser beam (14), the shaping and homogenization method of the PLD light beam (2) using a square-core multimode fiber (7) allows to control this important parameter in the OR-PAM scanner by selecting the fiber core (10) size and the magnification factor of the light direction module (13). Due to the nearly constant fluence profile and steep flanks, a high-contrast illumination point is also obtained, as shown in Figure 4a. The final average fluence reached at the square focus (15) with a side of about 75 μm is about 160 mJ / cm. 2 which corresponds to a measured power of 90 W and an energy of the light beam (2) of 9 μJ (measured in air). This fluence level is more than three times the maximum allowed to avoid damage to human skin for the 905 nm PLD wavelength used in this example, but the excess energy of the light beam (2) will help to compensate for the energy losses typically present in practical applications.

[0056] Furthermore, as already mentioned above, an improvement in the lateral resolution can be achieved by spatial filtering with a small aperture (11) centrally placed in front of the exit side (7") of the fiber core (10) using the same light direction module (13). The focal spot size (15) is then multiplied by a magnification factor m dThe reason for this is that the light direction module (13) now focuses the outgoing beam (14) with the same shape as the aperture (11) in the image focal plane (18) as before. Figure 4b shows the cross-sectional profile of the focal spot (15) resulting from spatial filtering by a diaphragm (11) with a circular aperture of 20 μm, with an increment of m in the light direction module (13). d =f 2 / f 1 ≈0.5. Thus the diameter of the focal spot (15) is about 10 μm, as represented by the circular line around the central peak in Figure 4b. It should also be noted that the pixel size of the CCD camera, 4.5 μm, prevents a better definition of the focal spot (15).

[0057] Although the energy of the light beam (14) reaching the focal point (15) in Figure 4b is consequently reduced by spatial filtering with the smaller aperture (11), the fluence at the focal point (15) remains approximately constant due to the pre-homogenization of the cross section of the light beam (2) at the end of the fiber core (10). The lateral resolution in this example is approximately 160 mJ / cm2, the same as for the larger focal point (15) of the square-core (10) optical fiber (7) without spatial filtering. 2 , and thus produces the same acoustic pressure amplitude, with a focal diameter of at least 10 μm (15).

[0058] In addition to the described OR-PAM technique and the corresponding embodiments of the scanning device, this optical method of shaping and homogenizing a light beam (2) based on an optical fiber (7) with a polygonal core (10) can be applied to other photoacoustic imaging techniques such as acoustic resolution photoacoustic microscopy, photoacoustic tomography, etc., to obtain a field of uniform fluence illumination from an inhomogeneous and low quality light source (3).

[0059] All of these imaging techniques can therefore benefit from the higher optical pulse (2) repetition rates of semiconductor-based light sources (3) compared to the more powerful and expensive solid-state lasers typically used as light sources (3) for these techniques.

[0060] Acoustic Resolution In both photoacoustic microscopy and photoacoustic tomography techniques, the lateral and axial resolutions depend only on the central frequency and bandwidth of the UST (21). However, to obtain a good field of view, a larger focal area with uniform illumination is generally required. This can be achieved by using an optical fiber (7) with a larger polygonal cross section, so that a higher power light beam (2) from a PLD with a larger emitting area can be efficiently coupled into the fiber (7) to provide a very good illumination uniformity and the required average fluence level at the imaging focal plane (18), as shown in FIG. 1 for an embodiment of an OR-PAM scanner. Furthermore, a larger focal area (15) can be scanned through this imaging plane (18) using a motorized mirror (17), albeit using a focusing lens (16) with a wider effective aperture and focal length.

[0061] In particular, for photoacoustic tomography scanners, image reconstruction is typically performed using a phased array ultrasound transducer (21), which is composed of multiple ultrasound detector elements, capable of capturing photoacoustic waves (20) at various viewing angles relative to the optically illuminated field of view. The phased array detector elements may be distributed around the light beam (2) in an annular or hemispherical geometry, similar to the single element annular transducer (21) in the light direction module (13) of the OR-PAM scanner shown in FIG.

Claims

1. An optical device (1) for shaping, homogenizing, guiding and focusing a pulsed light beam (2) onto a sample (19) of material, comprising: a light source (3) of a light beam or light pulses (2); a multimode optical fiber (7) having a polygonal cross-section core (10) with an optical inlet side (7') and an optical outlet side (7"); an optical coupling module (5) configured to guide the light beam or light pulse (2) from said light source (3) to the entrance side (7') of said optical fiber (7); a light direction module (13) configured to guide and focus the shaped homogenized light beam or light pulse (14) from the exit side (7″) of said optical fiber (7) onto a sample (19).

2. 2. The device (1) according to claim 1, comprising a surface (12) with an opening (11) arranged on the exit side (7'') of the optical fiber (7).

3. 3. The device (1) according to claim 2, wherein the aperture (11) is substantially centrally positioned relative to the core (10) of the fiber.

4. 3. The device (1) according to claim 2, wherein the size of the opening (11) is mechanically adjustable.

5. 10. The apparatus of claim 1, wherein the light source (3) comprises a pulsed laser diode, a light emitting diode (LED), or a stacked array thereof.

6. 2. The device of claim 1, wherein the multimode optical fiber (7) having a polygonal cross-section core (10) has a high numerical aperture.

7. 2. The device according to claim 1, wherein the light coupling module (5) and / or the light direction module (13) comprises two movable plano-convex aspherical lenses, one for collimation (27, 29) and the other for focusing (28, 16).

8. 2. The device of claim 1, wherein the light direction module (13) comprises one or more motorized scanning mirrors (17).

9. 1. A system for producing a photoacoustic image of a sample of material (19), comprising: An optical device (1) according to any one of claims 1 to 8, a transducer (21) for receiving acoustic waves (20), the transducer (21) being configured to generate an electrical signal from the reception of said acoustic waves (20); - A system comprising a unit for acquiring and processing the electrical signals generated by the transducer (21), the unit including software and / or hardware means configured to generate photoacoustic image data from the processed electrical signals.

10. 10. The system according to claim 9, wherein the ultrasonic transducer (21) is ring-shaped and coaxially surrounds the optical axis of the light beam (2).

11. 10. The system of claim 9, wherein the signal acquisition and processing unit comprises a preamplifier (22), an analog-to-digital converter (23), and / or a computer (24).

12. 10. The system of claim 9, wherein the signal acquisition and processing unit is connected to a screen or display (25) for real-time display of the generated optoacoustic image data.

13. Use of the device according to any of claims 1 to 8 for generating photoacoustic images in optical resolution photoacoustic microscopy applications, acoustic resolution photoacoustic microscopy applications and / or photoacoustic tomography applications.

14. A method for generating an optoacoustic image of a sample (19) of material, characterized in that the system according to claim 9 is operated to carry out the following steps: a) placing a light source (3) configured to emit a light beam or light pulse (2) with a duration in the nanosecond range and a frequency in the visible or near-infrared band at a distance from the sample (19) that is equal to or shorter than the length of the optical fiber (7); b) guiding the light beam or light pulse (2) coming from the light source (3) to the entrance side (7') of a multimode optical fiber (7) with a polygonal core (10) by adjusting the positions of the elements constituting the optical coupling module (5); c) focusing said shaped and homogenized light beam or light pulse (14) at a point on the sample (19) by adjusting the position of the elements constituting the light direction module (13); d) placing an ultrasonic receiving transducer (21) on the surface of the sample (19); e) electronically receiving and processing the signals captured by said transducer (21) using a signal acquisition and processing unit to obtain a photoacoustic image.

15. Steps c through e are repeated at least once; Each time, the position of the focal point (15) of the collecting lens (16) is changed by a second adjustment of the positions of the elements constituting the light direction module (13), 15. The method of claim 14, wherein a subsequent step of generating pixels corresponding to different points of the sample (19) in a plane perpendicular to the direction of the beam (2) and assembling the pixels is performed according to the irradiation position of the sample (19) corresponding to each of them, thereby obtaining a two-dimensional photoacoustic image of the sample (19).

16. The method is repeated at least once, Each time, the position of the focusing lens (16) is changed relative to the ultrasonic transducer (21) along the direction of the beam (2), 15. The method according to claim 14, wherein a subsequent step of generating photoacoustic images of different planes of the sample (19) longitudinally relative to the direction of the beam (2) and assembling said images is performed according to the irradiation positions of the sample (19) corresponding to each of them, to obtain a three-dimensional image of the sample (19).

Images