WO2021009205A2

WO2021009205A2 - Method and system for electromagnetic spectroscopy of a sample

Info

Publication number: WO2021009205A2
Application number: PCT/EP2020/069962
Authority: WO
Inventors: Riccardo CICCHI; Francesco Saverio Pavone; João Luis DE JESUS REIS LAGARTO
Original assignee: Consiglio Nazionale Delle Ricerche; L4T-Light4Tech Srl
Priority date: 2019-07-16
Filing date: 2020-07-15
Publication date: 2021-01-21
Also published as: WO2021009205A3; IT201900011904A1

Abstract

The invention relates to a method for electromagnetic spectroscopy of a sample, the method comprising: - Delivering an excitation beam of electromagnetic radiation to a first portion of a sample to obtain emission, or elastic or inelastic scattering of photons, said photons having a wavelength in the visible range, from molecules forming the sample; - Illuminating the first portion of the sample with a visible light, said visible light being switched on and off at a first frequency, so that illuminated time intervals and non-illuminated time intervals of the sample are defined; - Detecting said emitted, or elastically or inelastically scattered photons having a wavelength in the visible range during said non-illuminated time intervals; and - Obtaining spectroscopy information from said detected emitted, or elastically or inelastically scattered photons. It also relates to a system for electromagnetic spectroscopy of a sample.

Description

Method and system for electromagnetic spectroscopy of a sample

Technical field

The present invention relates to a method and system for electromagnetic spectroscopy of a sample under external illumination by visible light so that an operator can see the sample while the method or the system is in use.

Background art

Imaging and spectroscopy, for example fluorescence imaging and spectroscopy, have been increasingly exploited for sensitive and quantitative characterization of metabolic activity and structural alterations in biological tissues ex vivo and in vivo. Specifically, autofluorescence measurements are particularly attractive for clinical research, as they exploit the photo-physical properties of endogenous fluorophores (e.g. collagens, elastin, NAD(P)H, keratin) to provide label- free contrast and report structural and functional properties of biological tissues, thereby avoiding the use of potentially toxic exogenous contrast agents. The clinical potential and feasibility of autofluorescence measurements for in vivo clinical diagnosis have been demonstrated in a number of studies, either employing steady-state, and/or time-resolved autofluorescence measurements. Steady-state autofluorescence intensity or emission spectroscopy has been widely employed in clinical diagnostic research. However, intensity measurements alone are sensitive to intensity artefacts and so are difficult to compare between samples and patients. Spectrally-resolved ratiometric measurements aim to provide quantitative readouts but their discrimination is limited by the broad overlapping emission spectra of many endogenous fluorophores. Time-resolved measurements can increase the specificity of the fluorescence acquisition by resolving the fluorescence decay dynamics and thus provide means to resolve endogenous fluorophores with overlapping emission spectra but different fluorescence lifetimes. Fluorescence lifetime measurements are relatively insensitive to intensity artefacts that affect other techniques. Hence, when combined with fibre-based instruments, fluorescence lifetime measurements of endogenous fluorophores could enhance clinical diagnostics and provide guidance in surgical procedures.

Time-resolved fluorescence measurements can be employed in time- and frequency-domain. The time-domain technique, employing time-correlated single photon counting (TCSPC), is typically regarded as the "gold-standard" method for fluorescence lifetime measurements, given its high temporal resolution, high sensitivity and dynamic range compared to other methods. TCSPC has been widely used in clinical research in both imaging and single point platforms. In imaging systems, TCSPC measurements are typically slow and impractical for real-time acquisitions, given that fluorescence decays need to be acquired for every single pixel. Alternatively, single point instruments making use of fibre optics can integrate the fluorescence signal from a larger sample volume and therefore gain on acquisition speed when compared to imaging platforms. A major drawback of single-point instrumentation refers to the lack of spatial information that is often required in clinical applications, i.e. to identify diseased tissue margins. Irrespective of whether it is implemented in an imaging or single point platform, an additional major limitation of TCSPC refers to its inability to distinguish fluorescence photons from background photons (e.g. from ambient room light or direct bright-field illumination), which makes TCSPC measurements impractical under bright background illumination. For this reason, most studies employing TCSPC detection are realized in pre-clinical settings, where room lights can be turned off and any other sources of background light can be more carefully controlled. The sensitivity of TCSPC to background light is perhaps the greatest hindrance to the adoption of this technique for guiding clinicians during medical procedures, since turning off all sources of background light is either impossible or, in the limit, a major disruption to the workflow and thus a potential hazard, which ultimately limits the widespread deployment.

This limitation, i.e. the inability to distinguish fluorescence photons emitted by the sample from photons produced by external light sources used for sample or room illumination, is not specific to

TCSPC alone, but may also affect other steady-state and time-resolved measurement techniques that employ photon detection in the visible range of the electromagnetic spectrum. This is the case of steady-state fluorescence spectral measurements employing a spectrometer or time-resolved fluorescence measurements via the time-gating technique.

Summary of the invention

An aim of the invention is to provide a method and system to obtain spectroscopy information from a sample while the sample is illuminated with visible light. Preferably, but not necessarily, the sample and the spectroscopy information derived from the sample are displayed on the same image that can be visualized and interpreted in real time.

According to a first aspect, the invention relates to a method for electromagnetic spectroscopy of a sample, the method comprising:

Delivering an excitation beam of electromagnetic radiation having a wavelength between 300 nm and 700 nm to a first portion of a sample to obtain emission, and/or scattering of photons, said photons having a wavelength in the visible range, from molecules forming the sample;

Illuminating the first portion of the sample with a visible light, said visible light being different and distinct from said excitation beam, said visible light being switched on and off at a first frequency, so that illuminated time intervals and non-illuminated time intervals of the sample are defined;

Detecting said emitted and/or scattered photons having a wavelength in the visible range during said non-illuminated time intervals; and Obtaining spectroscopy information from said detected emitted and/or scattered photons.

According to a second aspect, the invention relates to a system for obtaining electromagnetic spectroscopy information of a sample, the system comprising: - an excitation beam source to emit beam of electromagnetic radiation towards a first portion of a sample, said excitation beam source emitting a beam having a wavelength between 300 nm and 700 nm; a visible light source apt to illuminate the sample with visible light, the visible light source being different and distinct from the excitation beam source; - a switching element of the visible light source apt to switch on and off the light source at a first frequency, so that illuminated time intervals and non-illuminated time intervals of the sample are defined; a photon detector to detect emitted and/or scattered photons from the sample due to the excitation beam, said photon detector having a detection wavelength window in the visible range, and to emit a corresponding signal; and a control unit to activate said detector to detect emitted and/or scattered photons from the sample only during said non-illuminated time intervals or to activate a shutter to prevent emitted and/or scattered photons from reaching the detector during said illuminated time intervals, and to obtaining spectroscopy information from said signal function of the detected photons.

Spectroscopy concerns the investigation and measurement of the interaction of electromagnetic radiation with matter. The types of spectroscopy are categorized according to the type of radiation involved in the interaction. These include but are not limited to ultraviolet/visible spectroscopy, infrared spectroscopy or nuclear magnetic resonance spectroscopy.

In many applications, the spectrum is determined by measuring changes in the intensity or frequency of this energy. The types of energy studied include electromagnetic radiation and the corresponding spectroscopy is called electromagnetic spectroscopy. It also includes other radiations as well (such as particles or acoustic waves).

Techniques that employ electromagnetic radiation are typically classified by the wavelength region of the spectrum.

Further, spectroscopy can be classified as emission spectroscopy or scattering spectroscopy. Emission is the process by which a higher energy quantum mechanical state of a molecule becomes converted to a lower one through the emission of a photon, resulting in the production of light. The frequency of light emitted is a function of the energy of the transition. Since energy must be conserved, the energy difference between the two states equals the energy carried off by the photon. The energy states of the transitions can lead to emissions over a very large range of frequencies. For example, visible light is emitted by the coupling of electronic states in atoms and molecules (then the phenomenon is called fluorescence or phosphorescence). An excitation beam can in this case cause the emission of photons by the sample.

The photons may not only be emitted by the sample, i.e. the invention does not relate only to emission spectroscopy. Scattering spectroscopy can be considered as part of the invention as well. In this case, the excitation beam includes photons in the visible range that are scattered by the molecules in the sample. These scattered photons, still having a wavelength in the visible range, are detected by the photon detector and also from these photons spectroscopy information can be obtained. In other words, the invention relates to any spectroscopic spectrum that is obtainable from a sample after directing an excitation beam towards it and detecting the resulting photons from the sample, the photons being in the visible range.

In the present invention therefore, in order to study a sample, and obtain spectroscopy information from the same, the sample is irradiated with electromagnetic radiation, called excitation beam, having a given wavelength (or frequency). This wavelength may be varied within a tuning range. The selected wavelength(s) of the electromagnetic radiation depends on the material in which the sample is made and the states of the molecules that it is desired to excite. For example, this wavelength of the excitation beam could be between 300 nm and 700 nm. The excitation beam in this case has preferably a wavelength in the ultraviolet/blue range.

In the present invention, this excitation beam causes the sample either to emit photons, or the photons of the excitation beam are scattered by the sample. The photons which are emitted and/or scattered by the sample have a wavelength within the visible range.

In this text, with the term "scattering", both elastic and inelastic scattering are meant. Elastic scattering is a form of scattering where the energy of a photon is conserved, but its direction of propagation is modified. Inelastic scattering is a scattering process in which the energy of an incident photon is not conserved (in contrast to elastic scattering). In the same way, with the adjective "scattered", such as scattered photons, both inelastically and elastically scattered photons are meant. Therefore, unless otherwise specified, the scattering phenomenon discussed herewith encompasses both the elastic and inelastic one.

The visible spectrum is the portion of the electromagnetic spectrum that is visible to the human eye. Electromagnetic radiation in this range of wavelengths is called visible light. A typical human eye will respond to wavelengths from about 380 to 780 nanometres. In terms of frequency, this corresponds to a band of about 385 - 790 THz. In the following, with "visible range" of electromagnetic radiation, the above-mentioned ranges are meant.

This does not mean that all the photons emitted or scattered by the sample due to the excitation beam must have a wavelength within the visible range. It is enough that just some of the emitted and/or scattered photons have a wavelength within such range. Further, the emitted and/or scattered photons may have a plurality of different wavelengths, that is, they may not have all the same wavelength. Thus, the emitted and/or scattered photons may define an emission band, defined by their range of wavelengths. The photons which are used to obtain spectroscopic information in this invention are those in the visible range. The excitation beam might be for example produced by a laser. Preferably, the excitation beam is substantially a monochromatic beam. Preferably, the excitation beam is a pulsed beam, that is, the beam includes a plurality of pulses emitted with a given frequency. Preferably, the given frequency is within the M Hz range, for example from 1 MHz to 160 MHz, more preferably 1 - 100 Mhz. Preferably, the duty cycle of the pulsed beam is below 50%. The excitation beam is delivered to a specific portion of the sample, called first portion. Preferably, this portion is smaller than the whole sample. Preferably, it is much smaller than the whole sample. For example, the dimension of the first portion is between 0.1 mm and 3 mm (with "dimension of the first portion" the largest dimension of the illuminated first portion is meant, for example the dimension of the first portion is the diameter of the first portion in case of a substantially circular beam). The first portion is thus a portion of the surface of the sample.

This first portion specifies a surface portion of the sample but at the same time it also defines a volume portion of the same, specifically that volume portion which is affected by the excitation beam. The photons which are emitted and/or scattered due to the excitation beam delivered on the first portion are the photons giving the spectroscopic information of interest. For each time interval, a single first portion of the sample is thus considered. This way of considering a specific single volume of the sample for each measurement is also referred in the literature as "single point measurement". Single point measurements refer to measurements where the spectroscopic information is gathered from a single "point" (which can be thought of as a single pixel) at a time, the above mentioned volume portion of the sample. Because data are captured from a volume, only one location is measured at a given time, and thus single-point measurements lack spatial resolution. Single point measurements are typically carried out using fibre-optic probes.

This contrasts with imaging measurements, where information is captured from various independent points "simultaneously".

Further, in order for a user to see the sample during spectroscopy measurements, the sample is globally illuminated by a given light source, which emits electromagnetic radiation within the visible range. This light source could be for example a light emitting diode (LED). This light source is distinct and separated from the electromagnetic excitation beam. The excitation beam creates the photons to be detected, while the visible light illuminates the sample to allow a user to see it with his/her eyes. Indeed, also the visible light may excite photons in the sample, however those photons are not detected according to the method and apparatus of the invention, as detailed below. In other words, two separate and distinct electromagnetic radiation beams impinge on the sample: the excitation beam and the visible light beam. More beams can be used, however in the method of the invention there should be at least two beams, distinct and separated from each other.

Preferably, the visible light source may include a white LEDs, although LEDs of any wavelength can be used. Alternatively, standard "room light" can be used, which is modulated at 50-60 Hz (depending on the country), and it is then preferably synchronized with the spectroscopic measurement, as detailed below.

Furthermore, more than one visible light source can be used. For example, the visible light source may include two or more white LEDs, to increase intensity; one red LED, one blue LED and one green LED that illuminate the sample simultaneously to generate white illumination. Obviously, more than one LED per colour can be used as well.

The visible light source however does not illuminate the sample continuously, but it is switched on and off at a given first frequency. This first frequency might be constant or might change with time. The sample is therefore illuminated by this visible light in certain time intervals, when the visible light source is switched on, and not illuminated in other time intervals, when the visible light source is switched off. Only the first portion, and not the whole sample, need to be illuminated by the visible light source. Preferably, the whole sample is illuminated by the visible light source.

Preferably the non-illuminated time intervals are longer than the illuminated time intervals.

Preferably, the first frequency is in the range 10 Hz - 1 kHz, more preferably in the range 40Hz - 60Hz. Preferably, the first frequency is above or equal to 40 Hz. The frequency is chosen so that the light perceived by the user is comfortable (for example, it is perceived as continuous and not pulsed). In addition, the off time should be "long enough" to allow the desired measurements using a single photon detector.

As for duty cycle of the light emitted by the visible light source, i.e. the fraction of time during which the visible light source is on within one period, is preferably intrinsically related to the spectroscopic acquisition time. Preferably, the duty cycle is set as low as possible (to maximize acquisition time during the off cycle), as long as the sample is clearly visible to the operator. For example, the visible light source can be switched on for only 500 ps (2.5% at 50 Hz) as long as this provides sufficient illumination of the sample for the user to clearly see it. However, in a different embodiment, the opposite is also possible. For example, if the duration of the on time interval of the visible light source is of about 18 ms in 20 ms cycles (90% at 50Hz), this may be acceptable as long as the sample is clearly visible to the operator and there are enough photons emitted and/or scattered by the sample and collected during the off period of the visible light source. The invention further provides for a step in which the photons emitted and/or scattered by the sample due to the excitation beam are detected. The detection is realized for example by a photon detector, such as a photomultiplier tube (PMT), single photon avalanche diode (SPAD), hybrid detector or a charge-coupled device (CCD) camera. However, this detection occurs only when the visible light source is off, that is, in these time intervals when the visible light source is not illuminating the sample (in the not-illuminated time intervals). In this way, the collected photons are only due to the sample's emission and/or scattering caused by the excitation beam, and not to the visible light source used to make the sample visible to the user. Therefore, the detection of emitted and/or scattered photons occurs only during the non-illuminated time intervals and not during the illuminated time intervals. Alternatively, the emitted, or scattered photons are continuously detected during the illuminated and non-illuminated time intervals, but photons arriving during the illuminated time-interval are not considered in the subsequent processing step, that is, no signals derived from photons detected during the illuminated time intervals are used to obtain spectroscopic information. Alternatively, the detector is protected during the illuminated time interval using for example a mechanical shutter, so that no photons can reach the detector during the illumination by the visible light source.

The detected photons are used to obtain spectroscopic information, i.e. data, out of a signal generated by their detection. For example, their fluorescence lifetime could be calculated, that is, the fluorescence intensity decay information that is derived from the detected photons. In order to obtain the spectroscopic information from the detected photons, the signal derived from the detection of the photons, e.g. the signal outputted by the photon detector, is processed using suitable algorithms and techniques in order to derive from it the required spectroscopic information.

Preferably, the signal derived from the detection of photons is processed during the non-illuminated time intervals. Thus, in each non-illuminated time interval, the following processes may take place: - detecting photons due to the sample emission and/or scattering during each of the non- illuminated time intervals;

generating a signal function of the detected emitted and/or scattered photons; analysing the generated signal to obtain the spectroscopic information.

The last sub-step, i.e. the analysis, may continue also during a subsequent illuminated time interval, but it ends before the next non-illuminated time interval, that is, for each non-illuminated time interval j and subsequent illuminated time interval k, the analysing step ends before the non- illuminated time interval j+1 begins.

Preferably, the detection, which means the time in which the photons are collected, or in other words, the sensing, takes place in a time interval longer than 0.5 ms. The sensing (detection) time is preferentially comprised between 5 ms and 20 ms. This may thus set a constraint on the length of the non-illuminated time interval because there is a "minimum time" needed to collect enough photons to obtain reliable spectroscopic data.

Thus, within a non-illuminated interval and subsequent illuminated time interval, but before the next non-illuminated time interval starts, all the information which are needed for obtaining the desired spectroscopic information are gathered and calculated. For example, possible spectroscopic information mean fluorescence lifetime, fluorescence lifetime components, fluorescence intensity, ratiometric fluorescence intensity, fluorescence spectrum, spectral features of fluorescence. In this way, the spectroscopic information can be visualized substantially in real-time as soon as they are processed.

This method and system of the invention guarantee that emitted and/or scattered photons due to the excitation beam and background photons (in this case coming from the visible light source) are temporally resolved and the signal obtained from the detected photons is free from bright background light, which makes spectroscopic measurements possible and practical under bright background conditions. This method and system can be therefore used also during surgical procedures when illumination of the sample is essential.

This is possible due to the synchronization of the detection of the emitted and/or scattered photons with an external visible light source. In particular, a pulsed visible light source is used to provide periodic illumination of the sample. The spectroscopic measurement (the detection of the photons) is out of phase with the visible light source illumination, e.g. it is triggered when the light source is turned off and is stopped before another "on" cycle. In this manner, spectroscopic phenomena and background photons can be temporally resolved and thus the corresponding signal from the photon detector is background-free, which makes spectroscopy measurements possible and practical under visible light background conditions. Furthermore, due to the "short times" needed to process and elaborate the spectroscopic signals, they can be immediately rendered available to an operator.

Preferably, the excitation beam is delivered by means of a fiber-optic probe.

To further increase the impact and applicability of this method and system in clinical settings, excitation light delivery and corresponding photon collection can be preferably realized with a fibre- optic probe using a single point approach. Spectroscopy maps can be created and displayed for single point measurements by superimposing a visible guiding beam with the excitation beam, which creates a visible reference onto the sample that can be imaged and processed in real time. According to the first or second aspect, the invention may include, as alternative or in combination, one or more of the following characteristics.

Preferably, said first frequency is equal to or higher than 10 Hz. Even more preferably, the first frequency is comprised between 40 Hz and 60 Hz. . Preferably, the first frequency is comprised between 10 Hz and 1 kHz. Preferably, the duty cycle is comprised between 2% - 90% of the first frequency.

The first frequency is preferably set above 40 Hz because it is preferred that this frequency is sufficiently high to avoid the stroboscopic effect of a blinking source and thus the light will appear to be continuous to the human eye. More preferably, the first frequency is set at about 50 Hz because this corresponds to the alternating current (AC) frequency in most countries. Furthermore, the frequency is preferably such that a "on-line" visualization of the spectroscopic data is possible (so it should be "fast"), however at the same time the non-illuminated times need to be long enough to allow a collection of a number of photons that is high enough for obtaining the desired data. The frequency used as first frequency is thus preferably a compromise among all these different needs. Preferably, the method comprises: detecting the emitted and/or scattered photons at a second frequency f2;

wherein the first frequency fl and second frequency f2 have the following relationship:

fi

f₂ =— , where k = 1, 2, ...

K

More preferably, the method comprises: generating the first frequency by a first clock;

- generating the second frequency by a second clock, wherein the second frequency is the frequency at which the emitted and/or scattered photons are detected; and phase-locking the first and second clock.

As said, the detection takes place only when the sample is not illuminated by the visible light. Thus, preferably also the detection is performed at a second frequency so that there are time intervals in which the detection (sensing) take place and time intervals in which there is no detection. The minimum duration of the time interval in which detection takes place is preferably 0.5 ms. The preferred duration of the time interval in which detection takes place is comprised between 5 ms and 20 ms.

Preferably the first (fl) and second frequency (f2) are the same. The second frequency (f2) can also be created by frequency-dividing the first frequency (fl) so that the following criteria is met:

When k = 1 the two frequencies are identical. Preferably, the first frequency is generated by a clock. Even more preferably, two phase-locked clocks are generated: the first clock is the clock that generates the first frequency of the visible light source. For example, a possible first frequency could be set to 50 Hz and 10% duty cycle (i.e. 2 ms). A second clock is set for the spectroscopy measurements, that is, for photon detection and processing of the resulting signal to obtain the desired spectroscopic information. The second clock has thus a second frequency, for example set equal to the first frequency (e.g. set to 50 Hz) and phase shifted relative to the first clock. As an example, the phase shift might be equal to 50° which is equivalent to a 2.8 ms delay. The phase shift is preferably present so that the detection of photons by the sensors of the system "waits" a little bit after the switching off of the visible light source, in order to minimize the detection of photons originating from the visible light source.

Preferably, also the analysis of the detected photons is made at the same second frequency. In other words, every time the photons are collected in a given time interval (at the second frequency), the signals obtained due to the collected photons are analyzed and the desired spectroscopic data is obtained. Thus, the desired spectroscopic information is obtained at the second frequency. In this way, the spectroscopic information can be obtained real time, and for example displayed to an operator.

Preferably, the time intervals in which the detection of the photons take place and the time intervals in which the spectroscopic information are obtained from the signals generated by the detected photons are phase-shifted. There is obviously a delay between the beginning of the photon collection and the elaboration of the generated signal by the photons. The first clock is preferably used to directly modulate the visible light source, so that this is turned on and off at the first frequency, in the above example it is turned on for 2 ms (duration of the illuminated time interval) in every 20 ms clock period (thus the non-illuminated time interval lasts 18 ms). The second clock is used to trigger the photon detection and also preferably the spectroscopic measurement, such as for example a fluorescence lifetime measurement. The acquisition (detection) of the photons may for example start on the rising edge of the second clock and run for a pre-defined period At, which preferably is set in order not to overlap with the following illuminated time interval of the visible light source. This strategy guarantees that the measured spectroscopic information, for example a fluorescence signal, is not contaminated by the background light produced by the visible light source, since the two signals are temporally shifted. Alternatively, the first frequency can be generated by a photodiode detecting periodic variations in room light intensity, instead of a clock. For example, a photodiode can be used to detect periodic variations in room light intensity and trigger the spectroscopy acquisition when the intensity is found below a pre-defined threshold. Preferably, also the excitation beam is a pulsed beam. More preferably, the excitation beam has a third frequency comprised between 1 MHz and 160 MHz, more preferably the third frequency is comprised between 10 MHz and 80 MHz. The laser frequency should be in principle as high as possible without being "too high", consistently with the typical time of fluorescence decay. Being the fluorescence decay of endogenous molecules in the 1-10 ns range, the maximum preferred frequency is of about 100 MHz.

Preferably, the duration of the illuminated time interval by the visible light is comprised between 0.5 ms - 50 ms, more preferably between 1 ms and 5 ms. These values are related to the first frequencies mentioned above. There is a maximum value for avoiding perceiving light blinking and a minimum value consistent with the fact that a certain amount of light is required for seeing the sample, as well as with the typical switching time for LEDs.

Preferably, the method includes: providing a camera; delivering a guiding beam of electromagnetic radiation detectable by said camera on a second portion of a sample, correlated to the first portion, to detect the location of the guiding beam; capturing an image of the sample using the camera.

Preferably, the system includes: a camera; - a guiding beam source of electromagnetic radiation having a wavelength detectable by said camera apt to deliver said guiding beam of electromagnetic radiation on a second portion of a sample, correlated to the first portion, to detect the location of the guiding beam; a camera control unit to control the camera in order to capture an image of the sample using the camera.

Depending on the spectroscopic information desired, the excitation beam may have a wavelength outside the visible spectrum. This means that the excitation beam delivered to the first portion of the sample may not be visible to the human eye and, therefore, the exact location of the excitation beam on the sample is not known by the user, i.e. it is not known where the first portion is within the sample. However, it is desired to have images, preferably in real time, of both the spectroscopic information and the portion of the sample where the information is coming from. For this purpose, spectroscopic information, such as fluorescence lifetime maps, can be created and displayed by superimposing a guiding beam with the excitation beam, which creates a reference onto the sample, reference visible to the human eye, to the camera or to both, that can be imaged and segmented in real time. Thus, on an image of the sample, both the first portion of the sample is visualized (being identified by the location of the guiding beam projected on the sample surface) and the spectroscopic information obtained detecting and analyzing the emitted and/or scattered photons from that first portion of the sample.

Preferably, the guiding beam is delivered by means of a fiber-optic probe.

The electromagnetic radiation emitted by the guiding beam is in a wavelength which can be detected by the camera. Preferably, but not necessarily, the wavelength of the guiding beam is within the visible range for the human eye as well. In principle, any wavelength can be used for the guiding beam, as long it is within the sensitive range of the camera. Preferably, the guiding beam wavelength is out of the spectral range of the emitted and/or scattered photons to be detected. A possible selection for a guiding beam source can be a 785 nm laser, since blood has low absorbance at this wavelength.

Preferably, the electromagnetic radiation of the guiding beam has a wavelength in the visible range. In this way, the operator can see where the guiding beam is located.

Preferably, the guiding beam is a separated and different beam from the excitation beam and from the visible light. There are therefore preferably three separated and different beam impinging on the sample: the excitation beam responsible of the detected photons, the visible light to illuminate the sample so that it is visible to an operator and the guiding beam to allow an operator (for example using the eyes or a camera sensitive to the wavelength of the guiding beam) to understand where the excitation beam is impinging on the sample (the excitation beam may be in a wavelength not visible to humans).

Preferably, the method includes: directing the excitation beam and/or the guiding beam onto the first and/or second portion of the sample by means of an optical fiber.

In order to provide real-time acquisition and feedback, preferably a single point approach is used. This strategy maximizes photon collection by integrating the spectroscopic signal from an excitation volume of the sample. On the other hand, a single point instrument lacks the spatial resolution, which is relevant in many clinical applications requiring fine margin identification such as tumor identification in resection surgery. In order to address this limitation, preferably three steps may be added to the method of the invention. First step, an additional light source, for example but not necessarily in the visible spectrum, emitting a guiding beam, for example a continuous-wave (CW) laser source, may be added. The guiding beam preferably is superimposed with the excitation beam, but it can also be not perfectly coaxial with the same. Preferably, the guiding beam illuminates the sample in a second portion. Preferably the second portion is superimposed to the first portion.

Preferably the first and second portion are substantially coincident, that is, the excitation beam and the guiding beam are delivered to the sample on the same surface portion. The guiding beam, if visible to the human eye, can provide a visual reference to guide the operator during measurements. In case the wavelength of the guiding beam is not visible to the operator, but it is within the sensitivity range of a camera, the camera can detect and record the position of the guiding beam and its position on the sample can be displayed on a display in such a way that it is visible to the human eye. The operator can still therefore have a visual reference looking at the display although the excitation beam is not visible. Further, the camera, such as, for example, a color camera, is added to the setup to image the sample from a fixed position. The camera allows recording the sampling point, that is the location of the surface of the sample impinged by the excitation beam (and guiding beam), which is the first and/or second portion of the sample, standing still, that is, without moving, the whole sample is within the camera's field of view. Alternatively, the camera could also move and a processing algorithm is implemented. This is implemented in case of a "large" sample, that exceeds the field of view of the camera.

The guiding beam could be also always illuminating the sample. For example, the electromagnetic radiation forming the guiding beam could have a visible wavelength that is out of the spectral range for fluorescence detection. This means that the photons of the guiding beam are not detected by the detector, only by the camera. Alternatively, the guiding beam could be present only when the excitation beam is present. Preferably, the first and the second portion substantially overlap. More preferably, the first and the second portion substantially coincide. More preferably, the method includes: moving the excitation beam over the sample so that the excitation beam is delivered to a different first portion of the sample; correspondingly moving the guiding beam. The recording can also be realized as the operator moves the excitation beam and/or the guiding beam through different regions (i.e. through different first and/or second portions) of the sample. The spatial location of the first and the second portion therefore may vary with time. Preferably, the camera follows the movements of the guiding and excitation beams. Alternatively, the camera always captures images of the whole sample, so that the first and second portions are always visible. In this way, the spectroscopic information can be gathered for different portions of the sample and the images can be visualized in real time.

Preferably, the method includes the step of: capturing an image of the sample only during said illuminated time intervals. Preferably, the image of the sample is captured only when the visible light source is switched on, and only for the time in which it stays switched on. When the visible light source is switched off, the camera is preferably not capturing any image. Preferably, the first clock used to modulate the visible light source is also used to trigger image acquisition, so that images are acquired during the "on" cycle of the visible light source, i.e. during the illuminated time intervals. In this way, for example by means of image processing, the position of the guiding beam can be located from each image and subsequently the location of the excitation beam can be extrapolated.

The images are preferably captured independently on the excitation and guiding beam.

The camera is driven preferably at the first frequency. A single frame includes the image acquired during the illuminated time interval, so that it includes the field of view and the guiding beam (which is preferably always on).

As an example, by realizing fast image processing, the guiding beam (e.g. a 785 nm beam) can be detected and segmented from the image, allowing tracking the position of the measurement over time and in real time, and thereby providing spatial resolution to a single point measurement. Given that both spectroscopic information, such as fluorescence lifetime, and image processing preferably occur in parallel, one can overlay the fluorescence lifetime information on top of the white light image and create augmented reality maps that provide a visual feedback of the measurements in real time.

Preferably, the step of obtaining spectroscopy information from said detected emitted photons includes: obtaining fluorescent spectroscopy information from said detected emitted photons.

Fluorescence spectroscopy (also known as fluorimetry or spectrofluorometry) is a type of electromagnetic spectroscopy that analyses fluorescence from a sample. Therefore, in this embodiment, the electromagnetic spectroscopy involved is fluorescence spectroscopy. The photons detected are photons emitted from the sample due to fluorescence.

This embodiment relates to the field of ultraviolet-visible spectroscopy, in particular, fluorescence spectroscopy. Fluorescence is the spontaneous emission of a photon when an electron in a molecule relaxes to its ground state following excitation to a high energy level. The wavelength of a fluorescence photon is determined by the energy difference between the two electronic states involved in the transition. The wavelengths at which a fluorophore may be excited and emit fluorescence light depend on the electronic configuration of the molecules, which, in turn, depends on the chemical structure of the molecule and on its local environment. The average time between excitation of a molecule and subsequent emission of a fluorescence photon is referred to as fluorescence lifetime and this is typically in the time scale between 10¹² and 10⁸ seconds.

In this embodiment, according to the invention, the time-resolved fluorescence measurement is out of phase with the visible light illumination, e.g. fluorescent spectroscopy measurements are triggered when the visible light source is turned off and are stopped before another "on" cycle.

Preferably, said spectroscopy information includes fluorescence intensity decay information, i.e. fluorescence lifetime information.

The fluorescence lifetime or time decay of a molecule is defined as the average time that a molecule remains in the excited state upon absorption of light prior to returning to the ground state by emitting a photon. The photons are those emitted by the sample and detected by the spectroscopic system.

Further, the present invention preferably relates to fluorescence emission spectroscopy, where the emission of electromagnetic radiation from a sample due to the absorption of a beam of electromagnetic radiation, called excitation beam, is studied. Fluorescence emission spectroscopy can be formally classified into two categories: steady-state and time-resolved.

Steady-state fluorescence spectroscopy refers to measurements of fluorescence intensity and/or spectrum. Steady-state fluorescence measurements are typically accomplished using a CW radiation source to excite the sample at constant intensity. For detection of the emitted fluorescence photons, typical setups use a dispersive element to separate out the fluorescence emission according to its wavelength and a CCD detector to measure the light intensity at different wavelengths. Alternative configurations can employ a set of dichroic mirrors and multiple single channel detectors to realize ratiometric intensity measurements.

Time-resolved fluorescence spectroscopy refers to measurements of the fluorescence intensity in time following excitation using a short light pulse, which is typically provided by a laser source. Time- resolved fluorescence aims to resolve the characteristic fluorescence intensity dynamics, i.e. fluorescence lifetime, and thus provide additional information to that available from steady-state measurements alone. Fluorescence lifetime measurement can be realized in time- or frequency- domain, depending on the fluorescence excitation and detection strategies. In time-domain techniques, the fluorescence signal is recorded in time following excitation with a short optical pulse. Frequency-domain techniques measure the harmonic response of the system, from which lifetime information can be derived from the difference in phase and amplitude between a periodically modulated excitation signal and the resulting demodulated fluorescence emission.

The present invention preferably relates to time-domain measurements. These methods include but are not limited to time-correlated single photon counting (TCSPC) and time-gating/binning.

For time-resolved fluorescence spectroscopy, measured fluorescence intensity decays can be analysed in real-time, i.e. immediately after each acquisition completes, preferably using the phasor method. The phasor method consists in the Fourier Transformation of the fluorescence decay to produce Real and Imaginary components that can be represented in a polar plot, as g and s phasor coordinates, respectively. Phase (r_phase) and modulation (r_mod) lifetimes can be calculated for each decay curve, as shown in Eqs. 2 and 3:

1 s

^Tphase - _{2nf g} (2)

where / is the laser excitation frequency. For single-exponential fluorescence species, the characteristic phasor will fall on the universal circle, defined as the semi-circle centered at (g = 0.5, s = 0) and radius 0.5. A mixture of two fluorescence species with distinct single-exponential characteristics will lie inside the universal circle, as a linear combination of each individual component. Hence, any combination of the two fluorophores falls along a line connecting their characteristic single-exponential phasors. Given that the phasor analysis is a fit-free method, it is far less computationally expensive than other fluorescence lifetime analysis techniques, such as multiexponential least squares fitting. Therefore, it offers a rapid characterization of the fluorescence decays that is suitable for real-time implementation, as is required in our application. Any of the spectroscopic parameters calculated from the phasor analysis can be used to generate fluorescence lifetime maps, which are displayed in real-time for visualization by the user.

A comprehensive description of the phasor approach to fluorescence lifetime data is provided in:

M. A. Digman, V. R. Caiolfa, M. Zamai, and E. Gratton, "The phasor approach to fluorescence lifetime imaging analysis.," Biophys. J., vol. 94, no. 2, pp. L14-6, Jan. 2008. H. E. Grecco, P. Roda-Navarro, and P. J. Verveer, "Global analysis of time correlated single photon counting FRET-FLIM data," Opt. Express, vol. 17, no. 8, p. 6493, Apr. 2009.

C. Stringari, J. L. Nourse, L. A Flanagan, and E. Gratton, "Phasor fluorescence lifetime microscopy of free and protein-bound NADFI reveals neural stem cell differentiation potential.," PLoS One, vol. 7, no. 11, p. e48014, Jan. 2012. In case of a multispectral configuration, i.e. multiple detectors are used to separate the fluorescence signal in different wavelength bands, the phasor approach can be applied to the signal detected in each channel. Spectroscopy information of any channel can be displayed in real-time. Further, the fluorescence intensity in a specific wavelength band can be calculated as a ratio of the total fluorescence signal, i.e. a sum of the fluorescence signal measured in all wavelength bands, to report ratiometric spectroscopic results.

In this embodiment, according to the invention, the time-resolved fluorescence measurement is out of phase with the visible light illumination, i.e. time-resolved fluorescence spectroscopy measurements are triggered when the visible light source is turned off and are stopped before another "on" cycle.

Preferably, the method includes the step of: generating an image of said sample; and - visualizing said image and said spectroscopy information in a display.

Preferably, the spectroscopic information is displayed in the image together with the first or second portion of the sample. In this way, online manipulation of the sample is possible, for example during surgery, because it is evident where the spectroscopic information is coming from.

Preferably, said sample is a part of a biological organism. The sample could be part of an animal or a plant. The organism can be living or dead. The organism may include an animal, which includes humans. The part which is used as a sample can be connective tissue, muscular tissue, nervous tissue, epithelial tissue, a portion of an organ, etc.

Preferably, the method includes: identifying the location of said first portion in the sample by identifying the location of the second portion illuminated by the guiding beam.

Preferably, the step of obtaining spectroscopy information from said detected emitted photons includes: obtaining spectroscopic information using a time-correlated single photon counting step or a time-gating step. TCSPC is the most common and well-established method for fluorescence lifetime measurements. It is based on the principle that, at sufficiently low detection rates, all individual photons can be detected together with their respective time of arrival, measured relative to a pulsed reference signal. TCSPC utilises ultrashort excitation pulses at high repetition rates (in the MHz regime), photon counting detectors, such as PMT, SPAD or hybrid detectors, and high-speed counting electronics to accurately measure the time of arrival of single photons. In TCSPC, the time of arrival of each detected photon is measured relative to the laser excitation pulse and recorded. Measurements are repeated for additional excitation pulses and when sufficient photon events are recorded, a histogram of the number of photons across all of the recorded time points can be generated. The measurement scheme of TCSPC imposes that photons are detected one at a time, at an absolute maximum rate of one photon per excitation period. In practice, however, the dead time of TCSPC circuitry is in the order of 100 ns and therefore it typically spans multiple excitation periods. Given that a large number of photons is typically required to produce a fluorescence decay and to achieve a certain accuracy in the lifetime estimate, TCSPC is often regarded as a slow technique.

An alternative approach to TCSPC measurements is based on gated detection of the fluorescence intensity signal at different time points relative to the excitation pulse. Typical time gating approaches utilise a single gate that samples the fluorescence intensity decay over a time interval of interest. The process is repeated at different delays relative to the excitation pulse, until a sufficient number of photons are detected. The resulting fluorescence decay curve yields as many points as the number of time intervals sampled. TCSPC is detailed for example in Becker, W. The bh TCSPC Handbook; 2008, printed by Becker & Hickl GmbH; further information can be obtained in https://www.becker-hickl.com/the-bh-tcspc- handbook/.

A comprehensive description of the time-gating technique for fluorescence lifetime measurements is provided in:

Mcginty, J.; Galletly, N.P.; Dunsby, C.; Munro, I.; Elson, D.S.; Requejo-isidro, J.; Cohen, P.; Ahmad, R.; Forsyth, A.; Thillainayagam, A. V; et al. Wide-field fluorescence lifetime imaging of cancer. Opt. Express 2010, 1, 124-135. McGinty, J.; Requejo-isidro, J.; Munro, I.; Talbot, C.B.; Kellett, P.A.; Hares, J.D.; Dunsby, C.; Neil, M.A.A.; French, P.M.W. Signal-to-noise characterization of time-gated intensifiers used for wide-field time-domain FLIM. J. Phys. D. Appl. Phys. 2009, 42, 135103.

Gerritsen, H.C.; Asselbergs, M. a H.; Agronskaia, a V; Van Sark, W.G.J.H.M. Fluorescence lifetime imaging in scanning microscopes: acquisition speed, photon economy and lifetime resolution. J. Microsc. 2002, 206, 218-24.

Preferably, the excitation beam includes a first excitation beam at a first excitation wavelength and a second excitation beam at a second excitation wavelength.

Preferably, the excitation beam source is adapted to emit a first beam of electromagnetic radiation at a first wavelength and a second beam of electromagnetic radiation at a second wavelength towards a first portion of a sample.

Preferably, the first electromagnetic beam and the second electromagnetic beam impinge the sample sequentially. Therefore, for each time instant, only one of the first excitation beam and second excitation beam impinges the sample.

The autofluorescence signal emanating from biological tissues is inherently complex owing to the multitude of contributing endogenous molecules that have significantly different excitation and emission spectra. In order to obtain signal from multiple molecules, it is common to realize systems with multiple excitation laser sources of different wavelength, which are capable of exciting different endogenous fluorophores. Therefore, in the invention more than a single excitation source can be selected.

Description of the drawings

The invention will be now better understood with non-limiting reference to the appended drawings, where:

Figure 1 is a diagram depicting optical instrumentation of a system for spectroscopy according to the present invention;

Figure 2 is a graph of the timing scheme for TCSPC measurements, and image acquisition and processing in the system of figure 1;

- Figure 3 is a block diagram of the software implementation;

Figures 4a-4c are three different graph showing: (a) Measured background levels at different TCSPC integration windows (At in Figure 2) when LED is off (squares, measurements realized with ambient room light only) and when LED is pulsing at 50 Hz, interleaved with TCSPC measurements (circles); (b) Percentage of background light (ambient light and pulsed LED light) in the measured fluorescence decay of a fluorescence standard for a constant photon count rate of approximately 5x10^s photons per second. For the measurements in (a) and (b) the LED intensity was kept constant throughout the measurements at a count rate of 2.5x10^s photons per second; (c) Fluorescence lifetime variation with intensity for POPOP. Fluorescence intensity was adjusted by varying the excitation power. Probe-to-target distance was maintained constant throughout the measurements. Background light from pulsed LED was maintained constant at 2xl0⁵ photons per second throughout the measurements. Measurements were realized with a TCSPC window of 15 ms. A total of 500 measurements were realized for each data point;

Figures 5a-5c show Fluorescence lifetime map of the internal structures of lamb kidney showing autofluorescence contrast (a) Raw white light image and (b) white light image augmented with fluorescence lifetime data (T_phase); (c) Variation of fluorescence lifetime (top) and total fluorescence intensity (bottom) with acquisition time. The fiber was moved at an average speed of 19.4 mm/s. Scale bar = 10 mm;

Figures 6a-6d show the effect of fiber scanning speed in the fluorescence lifetime acquisition (a) White light image of bovine heart tissue (b) Fluorescence lifetime profiles in the region indicated by the dashed line in panel (a) for a slow (solid line) and fast (dashed line) moving fiber. White arrow points towards increasing distance.

Fluorescence lifetime maps for slow and fast measurements are shown in panel (c) and (d) respectively. In (c), the fiber tip was moved at an average speed of 17.3 mm/s. In (d), the fiber tip was moved at an average speed of 45.1 mm/s. Scale bar = 10 mm;

Figures 7a-7f show fluorescence lifetime maps of porcine articular cartilage (sample 1) before (top row) and after (bottom row) treatment with bacterial collagenase. (a, d)

White light images, (b, e) fluorescence lifetime maps (T_phase) and (c, f) corresponding phasor maps. Scale bar = 10 mm;

Figures 8a-8b disclose (a) Phasor map of articular cartilage post-treatment (b) White light image showing phasor-based segmentation. Regions 1 & 2 represent phasors enclosed in the rectangles 1 & 2 in the phasor map, respectively;

Figures 9a-9c represent (a) Spectral phasor map and (b) fluorescence emission spectra of pure POPOP (dashed line) and FAD (solid line), and mixture solutions at different concentrations (). (c) Fluorescence lifetime phasor maps for all solutions. Each column depicts a different solution with measurements of pure FAD and POPOP presented on the far left and right columns, respectively, and the two mixture solutions on the centre.

A total of 50 measurements were realized for each solution. Rows show post-processing results for different spectral binning. Each point in the phasor cloud represents a fluorescence decay along the spectrum. For example, top row illustrates the results of binning the entire spectrum in a single channel, resulting in a maximum of 50 phasor points (from n = 50 measurements) for each solution. Bottom row shows the results of no spectral binning i.e. 64 spectral channels, resulting in a maximum of 3200 phasor points per solution. Fluorescence decays with less than 100 photon counts at the peak were removed;

- Figures lOa-lOf represent autofluorescence lifetime maps of cartilage pre- (top row) and post-treatment (bottom row) with bacterial collagenase. (a, f) White light images of the cartilage specimen (b, g) Fluorescence lifetime maps generated by binning the entire array into a single channel. Panels (c-e) and (h-j) show fluorescence lifetime maps for different wavelength bands: channel 1 (400 - 450 nm); channel 2 (450 - 515 nm); channel 3 (515 - 580 nm). Wavelengths bands were generated in post-processing by spectrally binning 15 columns of pixels. Scale bar = 10 mm;

Figures lla-llf represent normalized autofluorescence intensity maps of cartilage pre- (top row) and post-treatment (bottom row) with bacterial collagenase. (a, d) Channel 1 (400 - 450 nm); (b, e) channel 2 (450 - 515 nm); (c, f) channel 3 (515 - 580 nm). Wavelengths bands were generated in post-processing by spectrally binning 15 columns of pixels;

Figures 12a-12f are graphs of average (a, b) fluorescence emission spectra and (c-f) fluorescence lifetime in regions of interest (ROI), as illustrated in Figs. 10(a, f). In panels

(a) and (b) solid lines depict average spectra and shaded regions the corresponding standard deviation;

Figure 13 represents image processing steps to detect in real-time a guiding beam in the white light image captured by the color camera;

Figure 14 is a diagram depicting another embodiment of optical instrumentation of a system for spectroscopy according to the present invention; and

Figure 15 is a graph of the timing scheme for TCSPC measurements, and image acquisition and processing of the system of figure 14.

Detailed description of preferred embodiments of the invention

The system for spectroscopy of a sample according to the invention is globally indicated with 1 in figure 1.

The system 1 includes an excitation light source 2 emitting an excitation light beam of electromagnetic radiation, such as a pulsed laser diode. The excitation light beam is directed towards a sample 3 by means of a fibre optic probe 4 also part of the system 1.

The system 1 further includes a camera 5, which may be, for example, fixed and pointing at the sample 3 to record the measurements and to capture images of the sample 3.

The excitation beam on the sample produces the emission or scattering of photons within the excited volume (first portion) of the sample 3. A fibre optic probe 4a is used to collect the emitted or scattered photons from the sample 3 and guide them towards a detector 7. The detector 7 emits a signal function of the detected photons. A processing unit 20 is connected to the detector 7 to record the signal outputted by the latter and calculate spectroscopic information from the same.

Further, a second beam, emitted by a guiding beam source 6, such as a laser diode operating in continuous wave (CW) mode, is added to provide a visible reference in the white light images captured by the camera 5, as described below. The guiding beam is directed towards the sample 3 by means of an additional fibre 4b. To prevent the guiding beam to reach the detector 7, a short-pass filter (not visible in the drawings) is preferably added to the detection optical pathway. The second beam has a wavelength within the sensitivity range of camera 5.

The system 1 also includes a visible light source 9, such as white LED, that illuminates the sample to provide bright illumination of the field of view (FOV) of the camera 5 during spectroscopic measurements. The visible light source is actuated by a LED driver 12 and switched on and off at a given first frequency by a clock 13. A display 21 is also preferably comprised in system 1 to visualize the image captured by camera 5 and the spectroscopic information.

The excitation beam and the guiding beam are both delivered to sample 3, preferably in the same location. Preferably, the excitation beam is a pulsed beam and the guiding beam is a continuous wave beam. The light source 9 is also switched on and off to periodically illuminate the sample and the image acquisition by camera 5 is also taken with the same frequency (clock 13 is also connected to camera 5).

The excitation beam generates a plurality of photons emitted or scattered by the sample 3. The photons are collected by fibre 4a and detected by detector 7. The collection of photons by the detector 7, and the switching on and off of white light source 9 are synchronized with each other. Preferably, also image capturing of the camera 5 is synchronized with source 9 and detector 7.

Detector 7 emits a signal for each detected photon that is recorded by processing unit 20 and spectroscopic information are therefore obtained. The camera and visible source are preferably triggered using the same clock to guarantee that the sample is well illuminated when the camera is acquiring.

An additional embodiment of a system for spectroscopy of a sample according to the invention is globally indicated with 100 in figure 14. The system 1 includes an excitation light source, including in turn a first excitation source 2a, emitting an excitation light beam of electromagnetic radiation at a first wavelength and a second excitation source 2b emitting an excitation light beam of electromagnetic radiation at a second wavelength. The first and second excitation source may be pulsed laser diodes. The first and second excitation light beams are directed towards a sample 3 by means of a first and a second fibre optic probe 4e, 4f, a probe for each excitation source, also part of the system 1. The two beams can be also directed towards the sample 3 using the same optical fibre probe.

The system 1 further includes a camera 5, the camera being as described in figure 1.

The first and the second excitation beam on the sample produce the emission and/or scattering of photons within the excited volume (first portion) of the sample 3. A fibre optic probe 4a is used to collect the emitted and/or scattered photons from the sample 3 and guide them towards a first and a second detector 7a, 7b. The detectors 7a, 7b emit signals function of the detected photons. Each detector 7a, 7b is provided for a specific excitation source, for example the first detector detects photons emitted or scattered due to the impingement by the first excitation beam and the second detector detects photons emitted or scattered due to the impingement by the second excitation beam. Alternatively, the detectors 7a, 7b detect within a specific wavelength range and this may be independent of the excitation beams.

For example: the first excitation beam has a wavelength of 375 nm and the second excitation beam has a wavelength of 440 nm. The first detector 7a detects between 400 - 500 nm and the second detector 7b detects between 500 - 600 nm. When the first excitation beam is impinging, both detectors 7a, 7b can be turned on. When the second excitation beam is impinging, only detector 7b can be turned on, since detector 7a range overlaps with the second excitation beam.

The first and second excitation beam impinge the sample alternatively: either the first excitation source is on or the second excitation source is on. The first and second excitation beam do not impinge on the sample at the same time.

A processing unit 20 is connected to the detector 7a and 7b to record the signal outputted by the detectors and calculate spectroscopic information from the same, in particular two sets of spectroscopic information, the first spectroscopic information due to the photons excited by the first excitation source and the second spectroscopic information due to photons excited by the second excitation source.

Further, a third beam, emitted by a guiding beam source 6, such as a laser diode operating in continuous wave (CW) mode, is added to provide a visible reference in the white light images captured by the camera 5. There can be a guiding beam 6 for each excitation beam, but it is not necessary. The functioning of the guiding beam 6 is identical to the one described with reference to the embodiment 1 of figure 1.

The system 1 also includes a visible light source 9, such as white LED, that illuminates the sample to provide bright illumination of the field of view (FOV) of the camera 5 during spectroscopic measurements. The visible light source 9 is identical to the one described with reference to embodiment 1 of figure 1.

The first or the second excitation beam and the guiding beam are both delivered to sample 3, preferably in the same location. Preferably, the first or second excitation beam is a pulsed beam and the guiding beam is a continuous wave beam. The light source 9 is also switched on and off to periodically illuminate the sample and the image acquisition by camera 5 is also taken with the same frequency (clock 13 is also connected to camera 5).

The first and second excitation beams generate a plurality of photons emitted and/or scattered by the sample 3. The photons are collected by fibre 4a and detected by detector 7a and/or detector 7b. The collection of photons by the detector 7a or 7b, and the switching on and off of white light source 9 are synchronized with each other. Preferably, also image capturing of the camera 5 is synchronized with source 9 and detectors 7a or 7b.

Detector 7a or 7b emits a signal for each detected photon that is recorded by processing unit 20 and spectroscopic information are therefore obtained. The camera and visible source are preferably triggered using the same clock to guarantee that the sample is well illuminated when the camera is acquiring.

Example 1

The system 1 used in this first embodiment is the one depicted in Figure 1. Pulsed excitation light beam is provided by a 375 nm laser diode 2 (BDL-SMN-375, Becker & Hickl GmbH, Berlin, Germany) operated at 20 MHz and guided to the sample 3 via a single 100 pm core diameter fiber 4 integrated in a custom-made quadrifurcated fiber bundle (EMVision LLC, Loxahatchee, FL, USA). Fluorescence emanating from the sample 3 is collected by seven 300 pm core diameter fibers 4a and, at the detection end, a lens relay focused the fluorescence light onto the cathode of a hybrid detector 7 (HPM-100, Becker & Hickl GmbH). An emission filter 11 (FF01-470/28-25, Semrock, Rochester, NY, USA), indicated in figure 1 with the general term of "optical detection system", restricted the collection of the fluorescence signal to the band 470 ± 14 nm. The output of the detector 7 is connected to a TCSPC acquisition card 20 (SPC-730, Becker & Hickl GmbH) that provided temporal resolution to the fluorescence measurements. The temporal resolution of the TCSPC card was adjusted to 8 bit, thereby dividing the fluorescence decay in 256 temporal bins, which is equivalent to 195 ps per bin at 20 MHz.

In order to provide spatial resolution to the single-point TCSPC acquisition, the system 1 includes a USB color camera 5 (DFK 33UP1300, The Imaging Source, Bremen, Germany) that is fixed and pointing at the sample 3 to record the measurements. A second laser diode 6 (FC-785-350-MM2-PC- 0-RM, RGBLase, Fremont, CA, USA) with center wavelength at 785 nm and operating in CW mode was added to the setup to provide a visible reference in the white light images, as described below. To prevent 785 nm light to reach the detector 7, a 700 nm short-pass filter (not shown in the drawings - ET700SP, Chroma Technologies, Bellows Falls, VT, USA) was added to the emission path. Finally, the experimental setup also includes a white light emitting diode 9 (LED, MNWHL4, Thorlabs, Newton, NJ, USA) that was directed at the sample 3 to provide bright illumination of the FOV during fluorescence lifetime acquisitions, as illustrated in Figure 1 and described in detail below. For measurements, the LED was located 30 cm away from the sample. The LED 9 was connected to a current controller (LEDD1B, Thorlabs) with suitable interfaces for external modulation. System 1 operates according to the method of the invention. TCSPC measurements are synchronized with LED 9 to realize real-time fluorescence lifetime imaging under bright background conditions. To generate fluorescence lifetime maps in real-time using TCSPC detection under bright illumination of the sample 3, three processes are synchronized: 1) TCSPC acquisition and corresponding analysis of the fluorescence decay using TCSPC acquisition card 20 ; 2) color image acquisition from camera 5 and reference beam detection; 3) illumination of the FOV by LED 9 without interfering with the fluorescence decay acquisition. The white LED 9, camera 5 and TCSPC acquisition are synchronized by a Nl PCI-6221 (PCI-6221 DAQ, National Instruments) data acquisition board (DAQ), which corresponds to the timing unit 13 in figure 1. The camera 5 and LED 9 are triggered using the same clock 13 to guarantee that the sample is well illuminated when the camera 5 is acquiring. The TCSPC acquisition is triggered 0.8 ms after the LED signal is set to low, to ensure that the LED 9 is completely turned off when the TCSPC measurement is initiated.

The timing diagram of this implementation is shown in figure 2. The whole TCSPC acquisition and analysis loop runs independently from the image acquisition, processing and display loop, as depicted in Figure 3. This strategy optimizes computational resources by parallelizing the two processes, thus making sure that the timing requirements of the acquisition are met. The main application was implemented in LabVIEW (LabVIEW 2015, National Instruments, Austin, TX, USA) running in a 64-bit desktop computer equipped with an Intel Core i7-7700 CPU 3.6GHz CPU, 16 GB of RAM and 2 Tb hard disk drive.

Coming back to figure 2, as can be seen from the first and second curve from above, the LED illumination by LED 9 and the imagine acquisition by camera 5 takes place at the same time. Thus, in each period of 20 ms, the LED is on for 2 ms and in this "on" interval, the camera 5 acquires the image. In order to realize fluorescence lifetime measurements under bright illumination of the sample, the white LED 9 is preferably out of phase with the TCSPC acquisition (see figure 2, second curve from below, "spectroscopic acquisition"). The LED is directly modulated using a 50 Hz square wave provided by a PCI-6221 board. At this frequency, the stroboscopic effect caused by the blinking light source is not recognized by the human eye and the operator effectively perceives a continuous bright illumination of the FOV. In order to avoid the large background caused by the LED, TCSPC measurements are realized when the LED is off. The duty cycle of the LED can be kept to a minimum to maximize the TCSPC integration time At (see figure 2) and photon collection. For a LED pulsing at 50 Hz with 10% duty cycle (i.e. 2 ms on-time), in theory, the maximum integration time for a TCSPC measurement would be 18 ms, after which the LED is turned on again. There are, however, some practical considerations that limit the temporal range of the TCSPC acquisition. In one hand, it is necessary to account for the LED switching latency to guarantee that the TCSPC measurement is free from background light originated by the LED. This is accomplished by starting the TCSPC acquisition with a slight delay relative to the falling edge of the LED trigger signal - this was set to 0.8 ms in our setup (50° phase shift), which is sufficient to accommodate the LED fall time (< 100 ps) - and by selecting a conservative integration time so that the measurement is completed before the LED is turned on again. The period for TCSPC measurements At is further limited since time must be allocated for (1) preparation of the measurement as per hardware requirements and (2) data storage, processing and display (see in figure 2 the last curve at the bottom, "data analysis", which ends before the new spectroscopic measurement starts, and the middle curve, "image processing and display"). While the first is virtually instantaneous, data storage, processing and display could be time- and resource-consuming and thus a bottleneck of this method. Hence, to provide fast online feedback of the measurement, fluorescence decays are analyzed immediately after acquisition using the phasor approach, as described above in the summary of invention. A characteristic phasor is retrieved for every measurement and displayed in polar coordinates, providing real-time feedback of the measurement. Altogether, in order to meet timing requirements, a conservative photon integration time is chosen together with computational-intensive algorithms that parallelize data storage and processing.

Illumination of the sample is provided only by the pulsed white LED 9 and thus all measurements presented here are realized with room lights turned off. The intensity of LED was adjusted for each sample, considering two conditions: 1) it provides sufficient output so that the specimen is well illuminated and visible to the operator at the naked eye; 2) the intensity of the LED is controlled to avoid exposing the detector to intense light that could cause permanent damage.

The LED is hardware driven, i.e. the PCI-6221 DAQ 13 directly feeds a 50 Hz signal via a BNC cable to the LED driver 12, the TCSPC acquisition is triggered via software: a 50 Hz signal that is locked and delayed by 2.8 ms relative to the LED signal is read by a PCI-6221 (13) analog-to-digital converter

(ADC) and the TCSPC acquisition is triggered when a rising edge of this signal is detected. In consequence, any delays in the software due to e.g. data processing and storage, may eventually result in not detecting the following rising edge event and in missed acquisitions. To guarantee the sample is well-illuminated during image acquisition, the USB camera 5 is hardware triggered using a 50 Hz square wave that is in phase with the LED trigger signal (see figure 2). The exposure time of the camera was set to 1.93 ms, thereby maximizing illumination during the frame acquisition. The resolution of each frame was set to 640 ^c 480 pixels. For the CW source 6, a 785 nm laser diode is used. In principle, any wavelength can be used as long as it is within the visible range of the human eye, detectable by the camera 5, and out of the spectral detection of the fluorescence signal. Through fast image processing, the 785 nm beam can be detected and segmented from the white light image, which corresponds to an approximate position of the fluorescence measurements, since the fluorescence excitation source and the 785 nm guiding beams are superimposed at the sample plane. In brief, detection of a 785 nm beam in a white light image encompasses the following steps (see for reference figure 13):

1. The image is acquired in step IF by the camera 5;

2. Extraction of the red channel from the RGB image, creating a greyscale 8-bit image, takes place in step 2F;

3. Intensity threshold greyscale image to create a binary mask in step 3F; 4. Morphological erosion and closure on binary mask, to erode away the boundaries of the foreground object and remove noise in step 4F;

5. Morphological closure to fill holes inside object 6. Detection of the object contours in step 5F, the result of this step 5F is also shown in an enlarged view of the area of interest in figure 13;;

7. Fit circle to the detected contour and creation of a circular mask (step 6F);

For a fixed threshold, the diameter of the mask depends on various experimental parameters including FOV illumination, probe-to-target distance and tissue absorption and scattering properties.

For each frame where the reference beam is found, pixels within the circular mask are attributed the last lifetime value output by the phasor analysis. If a pixel is found within the circular mask in multiple acquisitions, its lifetime value is averaged out (see steps 7F and 8F of figure 13). This allows to produce a dynamic fluorescence lifetime map that can be superimposed with the raw white light image and updated at 50 Hz, thereby providing a visual feedback of the measurements in real time.

Fluorescence intensity decays were analyzed using the phasor method above described.

As described, each fluorescence decay is acquired from a group of pixels that is segmented to generate fluorescence lifetime maps. Therefore, the reciprocity between the fluorescence decay and phasor transformation determines that each point in the phasor cloud can also be traced back to a group of pixels in the fluorescence lifetime image.

The fluorescence lifetime so calculated (step 8F) is also superimposed to the image (step 9F).

The step 9F of fluorescence lifetime calculation is depicted in detail in figure 3 and its interaction with the image processing.

The above system and method have been applied to detect fluorescence information from animal tissues. Bovine heart and lamb kidney samples were acquired and kept at 4°C for a maximum of 12 hours before measurements. Porcine articular cartilage samples were obtained from trotter joints of freshly slaughtered pigs delivered from the local abattoir. The articular surface from metacarpophalangeal joints were exposed. Cartilage was left attached to the subchondral bone, which was cut distally using a hacksaw. Following extraction, cartilage pieces with approximately 3 ^c 3 x 3 cm were kept in PBS with 0.05% sodium azide for 24 hours to prevent bacterial growth. Samples were thoroughly washed in Phosphate Buffer Saline (PBS) and kept at -20°C for 24 hours before measurements.

Porcine articular cartilage samples (n = 2) were prepared for fluorescence lifetime measurements cartilage digestion. Localized lesions in the articular cartilage surface were induced by soaking filter paper (approximate dimensions 5 x 4 mm) in 250 mM bacterial collagenase. For control purposes, filter paper soaked in PBS was also placed in the articular surface. Treatment was applied for 5 hours at 37°C. Measurements were realized before and after treatment. The results presented below are relative to sample one only, similar results were obtained for sample two.

To demonstrate the feasibility of this method, it has been investigated whether background light originating from the LED was contaminating the TCSPC readout, for different integration windows Dί (see figure 2). For comparison, the background light produced by ambient room light only has been measured, i.e. the LED 9 was turned off during the entire acquisition. The results are presented in figure 4a. As expected, the background level originating only from ambient light increases linearly with integration time (see figure 4a, squares). When the LED 9 is turned on and pulsing at 50 Hz (circles), a linear increase in measured background up to an integration time of 15 ms is also observed. For integration times longer than 15 ms, a large increase in the number of photons acquired is measured, which suggests that light originating from the LED 9 is being detected by the TCSPC system 20. This increase can be explained by the timing constraints of the present method (see figure 2): if the TCSPC integration time is longer than 17.2 ms (which corresponds to the time difference between the start of the TCSPC measurement and the next LED cycle), it will likely overlap with the 2 ms LED illumination window. Delays in the software further decreases the maximum integration time range. Overall, these measurements demonstrate that for integration times up to 15 ms the LED produce a residual background level (less than 20 photons in 15 ms measurements) that does not affect the TCSPC readout. From 5 to 15 ms, an increase in background when the LED is turned on (circles) relative to ambient room light only measurements (squares) is observed. This could be attributed to LED background still being detected particularly at the edges of the TCSPC integration window and also to detector after pulsing: while the LED output is not significantly contaminating the TCSPC measurement, it is still reaching the detector and producing afterpulsing artifacts. To further investigate the viability of this method, the percentage of background originating from the pulsed LED 9 and ambient room light in the fluorescence decay of a fluorescence standard (see figure 4b) is measured. The background light was kept at a constant rate of 2.5x10^s photons per second (constant fraction discriminator (CFD) rate). When the fluorescence standard was excited, the CFD rate increased to approximately 5x10^s photons per second. The results in figure 4b demonstrate that the percentage of background in the fluorescence decay is residual (approximately 1%) for integration times up to 15 ms. As expected, for longer integration times the contribution of background increases to approximately 45% of the total measured signal due to the overlap of the TCSPC measurement with LED illumination, which is consistent with the increase measured in the CFD rate.

A common concern often associated with TCSPC measurements refers to the long integration times that are typically necessary to collect sufficient photons to accurately describe a fluorescence decay. From the results presented in figure 4(a, b), the maximum viable integration time for the present system is 15 ms. Hence, the ability of system 1 to measure the fluorescence lifetime of a reference fluorophore presenting single exponential characteristics (POPOP in ethanol) using a fixed integration time of 15 ms is verified. Measurements were realized for different laser excitation intensities to assess the ability of the system 1 to accurately measure the fluorescence decay of POPOP at different fluorescence intensities. The results presented in figure 4c show that at higher photon counts the measured fluorescence lifetimes are in close agreement to the reported value of 1.29 ns. As expected, the precision and accuracy of the measurements increase with photon counts. Yet, for fluorescence decays with approximately 500 photons we measured a lifetime of 1.32 ± 0.09 ns, which is still in excellent agreement with the expected value. Thus, a fixed integration time of 15 ms was used in all measurements reported below.

Following initial validation and characterization of the TCSPC acquisition strategy, label-free fluorescence lifetime maps of biological specimens out of single point measurements, in real-time are produced. Figures 5 a-c show the fluorescence lifetime measurement of a bisected lamb kidney. Lamb kidney is a convenient test sample for autofluorescence measurements given its anatomical heterogeneity that constitute a rich source of endogenous contrast. The fluorescence lifetime data (see figure 5b) is in general agreement with previous TCSPC measurements of lamb kidney. In particular, blood vessels (see white arrows in Figure 5a) present longer autofluorescence lifetimes relative to other regions, due to the higher elastin and collagen content and their characteristic long lifetimes. The measured fluorescence intensity varies significantly with an average of 8568 ± 7514 photons per each TCSPC measurement (see figure 5c, n = 7163 acquisitions).

Further, fresh bovine heart tissue was scanned at different speeds. Specifically, fluorescence lifetime maps were generated with average scan speeds of 17.4 mm/s (figure 6c) and 45.1 mm/s (figure 6d), corresponding to "slow" and "fast" acquisitions, respectively. The results demonstrate the similarity between measurements, which is further confirmed by plotting the measured fluorescence lifetime over the green dashed line (see figure 6(a-b)). The impact of scanning speed reflects essentially on the number of times each pixel is sampled: for lower scanning speeds, each pixel is sampled a greater number of times, thereby producing more consistent fluorescence lifetime results through averaging.

Consequently, measurements at slower scanning speeds produce smoother fluorescence lifetime maps, as demonstrated in figure 6(c-d).

To illustrate the potential to report intrinsic contrast in a clinically relevant application, porcine articular cartilage samples were mapped before and after localized treatment with bacterial collagenase to promote digestion (see figure 7a-f). The autofluorescence signal from articular cartilage emanates predominantly from collagen type II crosslinks. Previous studies demonstrated that fluorescence lifetime measurements are sensitive to enzymatic degradation of cartilage, either by cleavage of the collagen fibrils or depletion of proteoglycans. Before treatment (see figures 7(a- c)), an average lifetime of 6.89 ± 0.45 ns over the entire sample, for a total of 1528 TCSPC measurements and a period of 30.8 seconds, was measured, resulting in an average TCSPC rate of 49.6 Hz. Little variation of fluorescence lifetime over the entire sample (see figure 7b) was observed, which resulted in a compact phasor cloud, as shown in figure 7c. Treatment with 250 mM bacterial collagenase for 5 hours resulted in visible digestion of the articular cartilage surface, as shown infigure 7d, which translated in a drastic decrease in fluorescence lifetime in the affected area (see figure 7e). In the control region, a slight decrease in fluorescence lifetime was seen, which is in agreement with previous reports.

Decrease in fluorescence lifetime in the affected region generated a broader phasor cloud, with a significant spread towards shorter fluorescence lifetimes, i.e. clockwise rotation relative to the pre treatment phasor map (see figure 8a). The phasor cloud of the pre-treated sample falls within the region enclosed by the rectangle identified with numeral 2 in figure 8a, which suggests that these phasors originated from measurements in the region where treatment was not applied. Taking advantage of the reciprocity between the fluorescence decay measurements and corresponding phasor transformation, and pixel segmentation, phasors within rectangles 1 and 2 in figure 8a can be mapped and highlighted in the white light image, while excluding phasor points outside these regions. The results of this phasor-based segmentation are shown in figure 8b, where regions 1 and 2 represent phasors within the rectangles 1 and 2 in figure 8a, respectively. Region 1 overlaps to great extent with the region where treatment was applied, which further demonstrates the benefit of realizing phasor analysis. Given the simplicity of the phasor approach and its suitability for real-time implementation, a phasor-based segmentation that provides discrimination between normal and diseased tissues could also be implemented in real-time, thus increasing the impact for clinical applications.

Example 2

In this example, the system for fluorescence spectroscopy is the same as in figure 1 except for the type of detector 7 used. A SPAD array is used. In this second embodiment of system 1, fluorescence photons from the sample 3 were dispersed by a transmission grating 11 ("Optical detection system" of figure 1, GT50-06V, Thorlabs, Newton, NJ, USA) and the resulting spectrally-resolved fluorescence signal imaged with 1:1 magnification across the long axis of a SPAD array 7 (SPC3, MPD, Bolzano, Italy) consisting of 2048 pixels arranged in 32 rows by 64 columns. A 400 nm long-pass filter (FEL0400, Thorlabs) was added to the emission path to prevent excitation light reaching the detector. This optical configuration is simple and has a broad wavelength range (400 - 650 nm).

To generate autofluorescence lifetime and intensity maps out of single-point measurements, the same excitation beam and aiming beam of Example 1 are used. In brief, fluorescence measurements are realized by moving the fiber probe 4 freely across the sample 3. A second light source - a continuous-wave (CW) 785 nm laser diode 6 (FC-785-350-MM2-PC-0-RM, RGBLase, Fremont, CA, USA) - is used in simultaneous with the 375 nm excitation light to provide a visible reference to the measurements that can be recorded using a color camera (DFK 33UP1300, The Imaging Source,

Bremen, Germany). Rapid image acquisition and processing allows determination of the position of the 785 nm beam in real-time, which is approximated to the location of fluorescence excitation and collection. To prevent contamination of the fluorescence signal with 785 nm light, a 700 nm short- pass filter (ET700SP, Chroma Technologies, Bellows Falls, VT, USA) was also added to the emission path. Illumination of the field of view (FOV) was provided by a white light emitting diode 9 (LED, MNWFIL4, Thorlabs) modulated by a square wave at 50 Hz and with 10% duty cycle, i.e. 2 ms on cycle as in Example 1. Finally, triggering the color camera 5 simultaneously with the LED 9 will produce well- illuminated color images at 50 Hz (with ~2 ms exposure time), which can be used for determining the region of measurements in real-time, as described above. Given that camera and fluorescence acquisitions are synchronized, each fluorescence measurement is associated with a position of the fiber probe. In consequence, fluorescence lifetime and spectral maps can be generated and augmented on the white light image to provide visual feedback of the measurements in real-time. One particular highlight of this method is that spectral bands can be selected and tuned on-the-fly during processing, so that online feedback is provided for a specific region of the spectrum, and the spectral band can be changed if necessary.

The detector 7 used in this second embodiment is an array of 32 ^c 64 SPADs (4.8 ^c 9.6 mm) developed using 0.35 pm complementary metal-oxide semiconductor (CMOS) technology. Histogram of photon arrival times are generated independently in each pixel by means of time-gating, at a fixed repetition rate of 50 MHz, which was used for synchronization with the excitation laser. Since the integration time of this application is limited, in order to improve photon collection efficiency while maintaining reasonable sampling of the fluorescence intensity decays, the used gating strategy is to employ relatively long detection gates (4 ns) and 40 sampling points. The acquisition window was adjusted to 16 ns. All measurements presented were realized with these gate settings. The reference fluorophores are the following.

Stock solutions of flavin adenine nucleotide (FAD, F6625, Sigma-Aldrich, Saint Louis, MO, USA) and l,4-bis(5-phenyloxazol-2-yl) benzene (POPOP, P3754, Sigma-Aldrich) were prepared by dissolving the corresponding powder in 50 mL of purified water and ethanol, respectively, to achieve concentrations of 50 mM. Two mixture solutions were obtained from stock solutions with the approximate proportions 1:2 and 2:1 ([FAD]:[POPOP]).

The fluorescence decay characteristics for each set of segmented pixels were analyzed in real-time using the phasor approach, as in Example 1, using equations 2) and 3). Autofluorescence lifetime data presented here report r_phase measurements. Porcine articular cartilage samples were prepared as previously described in Example 1. Briefly, articular cartilage specimens (dimensions 3 x 3 x 3 cm) were obtained from metacarpophalangeal joints of freshly slaughtered pigs and kept in PBS with 0.05% sodium azide for 24 hours to prevent bacterial growth. Articular cartilage digestion was induced using filter paper soaked in 250 mM bacterial collagenase on the articular surface. Treatment was applied for 5 hours at 37°C. Fluorescence measurements were realized before and after treatment over the entire cartilage surface to demonstrate spatial variation of the fluorescence signal. Samples were thoroughly washed in Phosphate Buffer Saline (PBS) and kept at -20°C for 24 hours before measurements. A total of n = 2 samples were prepared and measured. The results presented below are relative to one sample only. Consistent results were obtained for the second sample. Fluorescence spectral and lifetime measurements of reference fluorophores are as follows.

The dynamic spectral lifetime resolution can be tuned according to the requirements of each application. In order to verify the impact of the spectral resolution in the fluorescence output, the fluorescence lifetime and spectral characteristics of pure FAD and POPOP are measured, and two mixture solutions (see Figure 9a-9d). For pure solutions of POPOP (see Figure 9(a, b) dashed line) and

FAD (see Figure 9(a, b) solid line), we measured maximum fluorescence emission at ~425 nm and 530 nm, respectively, which are in good agreement with previous reports. For the two mixture solutions, the fluorescence spectra show traces of both fluorophores in different proportions: when FAD is predominant (+FAD curve), we measured maximum emission at "'530 nm and a shoulder at 410-430 nm consistent with POPOP emission peak; when POPOP is the predominant fluorophore (+POPOP curve), maximum emission occurs at 420 nm and is accompanied by a shoulder at 510-520 nm, which is indicative of FAD fluorescence.

Fluorescence lifetime data for each solution are presented in Figure 9c, where each column depicts a different solution, with measurements of pure FAD and POPOP presented on the far left and right columns, respectively, and the two mixture solution on the center. To demonstrate how spectral resolution impacts the fluorescence readout, data were spectrally binned in post-processing to adjust the spectral resolution of the measurements, from its maximum (Figure 9c bottom row, 64 spectral channels, corresponding to 1 column of pixels per channel) to its minimum (Figure 9c top row, 1 spectral channel, corresponding to 64 columns of pixels merged into a single channel). Each point in the phasor cloud represents a fluorescence decay along the spectrum, i.e. for a defined spectral band. Consequently, the greatest the spectral resolution, the larger the number of phasors in each plot. For pure FAD and POPOP (far left and far right columns, respectively), compact phasor clouds independently of the spectral resolution are obtained. This indicates that the fluorescence lifetimes of both FAD and POPOP are independent of the wavelength. In turn, for mixture solutions, the fluorescence specificity increases with spectral resolution. This is most evident when POPOP is the dominant fluorophores (see Figure 9c, third column). Specifically, when data are merged into a single channel (Figure 9c, panel 3), the corresponding fluorescence decay will be a weighted average of the fluorescence decays across the spectrum. Since POPOP is predominant over FAD (see Figure 9b), the corresponding phasor will be closer to the pure POPOP phasor. If data are binned into four equal spectral channels (Figure 9c, panel 7), the specificity of the measurement increases and two clusters of phasors are now evident in the phasor map, corresponding to the FAD and POPOP spectrally- resolved fluorescence signals. A third cluster is also visible along the line connecting the pure populations, representing fluorescence decays within the spectral region where POPOP and FAD have competing contributions, i.e. from ~450 nm to 500 nm. As the spectral resolution increases (see Figure 9c, panels 11, 15 and 19), the number of phasor clusters along the line connecting the pure populations also increases and their distribution varies depending on the contribution of each fluorophore in each spectral band: fluorescence decays from short wavelength bands are shifted towards pure POPOP (to the right); in turn, fluorescence decays from long wavelength bands are closer to FAD (to the left). For maximum spectral resolution, i.e. 64 spectral channels, fluorescence decays in each channel have a slightly different contribution from each fluorophore, resulting in slightly different fluorescence lifetimes. Accordingly, the corresponding phasors occupy the entire region between FAD and POPOP pure populations, shifting from the pure POPOP towards FAD as the wavelength increases (Figure 9c, panel 23).

To demonstrate the implementation in clinically relevant biological tissue, the autofluorescence lifetime and spectral signatures of porcine articular cartilage has been measured. Measurements were realized before and after treatment of the articular surface with bacterial collagenase, to induce endogenous contrast derived from the localized digestion of collagen and consequent alteration of its fluorescence characteristics. The photon integration time for each autofluorescence acquisition was set to "'ll ms. For this temporal window, a negligible contribution from background light originating in the LED is present. For longer integration times, there is significant increase in the contribution of LED light to the autofluorescence measurement, which results from delays in data acquisition and processing that cause the fluorescence acquisition to overlap with the on cycle of the LED. This is essentially because the fluorescence acquisition is software triggered and thus more susceptible to delays and timing errors. For this reason, while fluorescence and image acquisitions are triggered at 50 Hz, some acquisition cycles are skipped, and the real acquisition rate can vary between 25 and 50 Hz depending on computer processor usage and memory consumption. Nevertheless, an acquisition rate of 25 Hz is still sufficient to provide feedback of the measurements in real-time.

Autofluorescence lifetime of porcine articular cartilage pre- and post-treatment with bacterial collagenase are presented in Figure lOa-lOf for the entire spectral range (panels b, g) and for distinct spectral bands: channel 1, 400 - 450 nm (panels c, h); channel 2, 450 - 515 nm (d, i); channel 3, 515 - 580 nm (e, j). Normalized autofluorescence intensity maps of porcine articular cartilage pre- and post-treatment are presented in Figure lla-llc: normalized fluorescence intensity maps of cartilage pre- (top row) and post-treatment (bottom row) with bacterial collagenase. (a, d) Channel 1 (400 - 450 nm); (b, e) channel 2 (450 - 515 nm); (c, f) channel 3 (515 - 580 nm). Wavelengths bands were generated in post-processing by spectrally binning 15 columns of pixels. Figures lla-c show steady-state fluorescence data rather than lifetime and shows additional spectroscopy information retrieved from the same method as Figure 10.

During acquisition, online feedback was provided for the entire spectral range. Before treatment (see Figure 10(a-e)), a small variation of the autofluorescence lifetime across the articular surface has been observed (see Figure 12(c-f)), possibly arising from regions with different load-bearing characteristics. In the region where treatment was applied (see Figure 10a, square region, n = 1088 pixels), pre-treatment average autofluorescence lifetimes of 3.44 ± 0.03 ns, 3.20 ± 0.04 ns and 2.89 ± 0.05 ns have been measured for the wavelengths bands of channels 1, 2 and 3, respectively, which indicate a slight dependence of the autofluorescence lifetime with wavelength. Merging the data in a single spectral band (Figure 10b and Figure 12c), an average lifetime of 3.20 ± 0.05 ns is obtained.

Treatment with bacterial collagenase resulted in visible digestion of the articular cartilage surface (see Figure lOf, square region), which, in turn, resulted in a remarkable decrease of autofluorescence lifetime in this region relative to non-digested cartilage. While a decrease in autofluorescence lifetime is visible in all spectral bands, this is most evident in channel 2 (see Figure lOh), which is coincident with the fluorescence emission peak of collagen (see Figure 12(a, b)). A slight decrease in the autofluorescence lifetime of the non-digested region is measured, although this is not as evident as in the region of collagen digestion. In the region of digestion (n = 1740 pixels), we measured autofluorescence lifetimes of 2.82 ± 0.06 ns, 2.51 ± 0.07 ns and 2.36 ± 0.07 for channels 1, 2 and 3, respectively. For the entire spectral range, an average lifetime of 2.42 ± 0.05 ns was measured.

With respect to autofluorescence intensity (figure 12), data suggest a red shift in the autofluorescence spectrum of the digested region relative to non-digested cartilage. In general, a lower absolute autofluorescence signal after treatment is measured, which generated noisier autofluorescence maps and spectra. Example 3

In this example, the system 100 of figure 14 is used.

Excitation with multiple laser sources, such as first excitation source 2a and second excitation source 2b, may require the implementation of a laser switching scheme via pulse-group multiplexing.

The system 100 in Figure 14 comprises two excitation laser sources 2a and 2b (375 nm and 440 nm), and two photon counting detectors (detector 7a and detector 7b) for acquisition of the fluorescence signal.

Further, the system 100 comprises a fibre-optic probe 4a, 4e, 4f, a white LED 9 to provide illumination of the sample 3 and a USB camera 5 to record the measurement. The LED 9 is pulsing at a first frequency equal to 50 Hz (20 ms periods) with 2 ms illumination intervals and 18 ms non illumination intervals, see Figure 15. The two excitation lasers (excitation source 2a and excitation source 2b) can be turned on and off sequentially at a frequency of 25 Hz, which is synchronized with the LED illumination at 50 Hz (as visible in figure 15). During the time each laser is turned on (15 ms in this example, see Figure 15), the pulsing rate for fluorescence excitation is maintained in the MHz range, for example 20 MHz. With this multiplexing scheme, autofluorescence from different molecules can be obtained quasi-simultaneously and still provide real-time multidimensional feedback of the measurements. The signal obtained from either of the lasers can be collected by detectors 7a and 7b, thereby providing additional spectral resolution to the fluorescence emission. Because the signals obtained with each laser 2a, 2b are temporally separated (only one laser beam is impinging the sample in a predetermined time interval), they can be discriminated and processed independently.

The implementation of a system 100 with multiple laser sources has the advantage of allowing collection of broader and more diverse autofluorescence signal. For example, excitation at 375 nm permits measurements of NAD(P)H fluorescence, while excitation at 440 nm permits measurements of FAD fluorescence. This is relevant because these two molecules are intimately involved in metabolic processes and thus their fluorescence characteristics can report functional alterations, particularly when analysed together. Typically, measurements of NAD(P)H and FAD are realized sequentially but are temporally separated by seconds or tens of seconds, depending on the optical setup. Therefore, in such measurements it may not always be possible to correlate the fluorescence signals of NAD(P)H and FAD. In the implementation of the invention, measurements are still sequential but are realized within millisecond intervals, which makes the correlation between

NAD(P)H and FAD signals more robust and reliable. This is particularly relevant for a fibre-optic implementation, where the position of the fibre in constantly changing during acquisition.

Claims

1. A method for electromagnetic spectroscopy of a sample (3), the method comprising

Delivering an excitation beam (2) of electromagnetic radiation having a wavelength between 300 nm and 700 nm to a first portion of a sample (3) to obtain emission and/or scattering of photons, said photons having a wavelength in the visible range, from molecules forming the sample;

Illuminating the first portion of the sample with a visible light (9), said visible light being different and distinct from said excitation beam, said visible light being switched on and off at a first frequency, so that illuminated time intervals and non-illuminated time intervals of the sample are defined;

Detecting said emitted and/or scattered photons having a wavelength in the visible range during said non-illuminated time intervals; and

Obtaining spectroscopy information from said detected emitted and/or scattered photons.

2. Method according to claim 1, wherein said first frequency is comprised between 10 Hz and 1 kHz.

3. Method according to claim 1 or 2, wherein the detecting step last longer than 0.5 ms.

4. Method according to claim 3, wherein the detecting step last a time interval comprised between 5 ms and 20 ms.

5. Method according to one or more of the preceding claims, comprising:

detecting the emitted and/or scattered photons at a second frequency f2;

6. Method according to one or more of the preceding claims, wherein the excitation beam has a third frequency.

7. Method according to claim 6, wherein the third frequency is comprised between 1 MHz and 160 M Hz.

8. Method according to one or more of the preceding claims, wherein the duration of the illuminated time interval by the visible light is comprised between 0.5 ms - 50 ms.

9. Method according to claim 5, wherein the step of obtaining spectroscopic information includes:

- obtaining spectroscopic information at the second frequency.

10. Method according to one or more of the preceding claims, including the step of:

providing a camera (5);

delivering a guiding beam (6) of electromagnetic radiation detectable by said camera on a second portion of a sample, correlated to the first portion, to detect the location of the guiding beam;

capturing an image of the sample using the camera (5).

11. Method according to claim 10, including:

moving the excitation beam (2) within the sample (3) so that the excitation beam is delivered to a different first portion of the sample;

correspondingly moving the guiding beam.

12. Method according to claim 10 or claim 11, wherein the first and the second portion substantially overlap.

13. Method according to one or more of claims 10 - 12, comprising the step of:

directing the guiding beam onto the second portion of the sample by means of an optical fiber.

14. Method according to one or more of the preceding claims, comprising:

directing the excitation beam onto the first portion of the sample by means of an optical fiber.

15. Method according to one or more of the preceding claims when dependent on claim 10, including:

identifying the location of said first portion in the sample by identifying the location of the second portion produced by the guiding beam.

16. Method according to one or more of the preceding claims when dependent on claim 10, including the step of:

capturing an image of the sample only during said illuminated time intervals.

17. Method according to one or more of the preceding claims, wherein the step of obtaining spectroscopy information from said detected emitted and/or scattered photons includes: obtaining fluorescence spectroscopy information from said detected emitted photons.

18. Method according to claim 16, wherein said spectroscopy information includes fluorescent intensity decay information.

19. Method according to one or more of the preceding claims, including the step of: generating an image of said sample (3); and visualizing said image and said spectroscopy information in a display (21).

20. Method according to one or more of the preceding claims, wherein said sample is a part of a biological body.

21. Method according to one or more of the preceding claims, wherein the step of obtaining spectroscopy information from said detected emitted photons includes: obtaining spectroscopic information using a time-correlated single photon counting step or a time gating step.

22. A system (1) for obtaining electromagnetic spectroscopy information of a sample (3), the system comprising: - an excitation beam source (2) to emit beam of electromagnetic radiation towards a first portion of a sample (3), said excitation beam source emitting a beam having a wavelength between 300 nm and 700 nm; a visible light source (9) apt to illuminate the sample with visible light, the visible light source being different and distinct from the excitation beam source; - a switching element (12, 13) of the visible light source (9) apt to switch on and off the visible light source at a first frequency, so that illuminated time intervals and non- illuminated time intervals of the sample are defined; a photon detector (7) to detect emitted and/or scattered photons from the sample due to the excitation beam, said photon detector (7) having a detection wavelength window in the visible range, and to generate a corresponding signal; a control unit (21) to activate said photon detector to detect emitted and/or scattered photons from the sample only during said non-illuminated time intervals or to activate a shutter to prevent emitted and/or scattered photons from reaching the detector during said illuminated time intervals, and to obtaining spectroscopy information from said signal function of the detected photons.

23. System according to claim 22, wherein said photon detector is a photon-counting detector.

24. System according to claim 22 or 23 including: a camera (5);

guiding beam source (6) of electromagnetic radiation having a wavelength detectable by said camera apt to produce said guiding beam of electromagnetic radiation on a second portion of a sample, correlated to the first portion, to detect the location of the aiming beam;

a camera control unit (13) to control the camera in order to capture an image of the sample using the camera.