WO2025073586A1 - Detector device and method for determining a substance concentration - Google Patents

Abstract

A detector device comprises at least one light source (10) and at least one detecting component, wherein the at least one detecting component within a housing is optically separated from the at least one light source (10). The detecting component comprises at least three detecting areas (20,21,22), each detecting area arranged at different distances from the at least one light source (10). Each area (20,21,22) is configured to detect a light component corresponding to emitted light scattered through the sample. A size of each detecting area (20,21,22) is dependent on the wavelength of the emitted light and the distance to the at least one light source (10), such that a signal level and/or signal-to-noise ratio of the detected light at each detecting area is in a predefined common interval. A control circuit (2) is coupled to the at least one light source (10) and at least one detecting component and configured to obtain least three signals from said detecting component obtained from at least three detecting areas. The control circuit (2) is configured to derive a deviation of the at least three obtained signals or a processed curve thereof from a reference axis, said reference axis in particular including two of the at least three different signals.

Description

DETECTOR DEVICE AND METHOD FOR DETERMINING A SUBSTANCE CONCENTRATION

The present application claims priority form German application DE 10 2023 127 038 . 9 dated October 4 , 2023 , the disclosure of which is incorporated herein by reference in its entirety . The present invention relates to a method for determining a substance concentration in a sample containing particles in a liquid, in particular glucose in blood plasma , wherein a refractive index of the liquid is dependent on a concentration of the substance dissolved therein . The invention also relates to a detector device .

BACKGROUND

The current standard for blood glucose measurement often uses an invasive technique , in which a small amount of blood is drawn, and subsequent electrochemical analysis is performed using a handheld device . This method is not suitable for continuous monitoring because for each measurement , the finger must be pricked to obtain a fresh blood sample .

A more recently developed technology uses a button that sits on the skin and misses interstitial fluid in parts of the subcutaneous adipose tissue with a small , needle-like sensor . However , the needle penetrates permanently into the skin and may constitute a possible infection location .

Besides these invasive methods , there are also non-invasive methods based on optical IR measurements or Raman spectroscopy . While in the first case , a suitable choice of emitter and detector leads to difficulties , an approach based on Raman spectroscopy is challenging due to the very poor signal-to-noise ratio .

In this respect , there is a need for a method that can detect a substance in a liquid in a simpler way and allows for a continuous measurement . SUMMARY OF THE INVENTION

This and other obj ects are addressed by the subj ect matter of the independent claims . Features and further aspects of the proposed principles are outlined in the dependent claims .

Complementing some other ideas that are based on an angle dependent scattering , the present idea relies on an evaluation of the gradient of the scattering signal at different distances , which has been found by the inventors to also depend on the blood glucose concentration .

Due to scattering and absorption of light , which is initially directed towards the skin and may subsequently propagate through portions of shallow- and/or deeper lying tissue , which in turn hosts a network of blood vessels , the amount of light that re-emerges at the s kin-to- ambient interface some distance from the entry point will vary according to the optical path length ( distance ) travelled in blood . The path length is referred to as "blood optical path length ( BOPL ) " . The BOPL is affected among other things by the beating of the heart , but essentially by the location at which the scattered light signal is measured and more precisely from the distance between the location of the incident light and the location of the scattered light .

It is known that signal levels due to diffuse reflection at a given distance from the illumination spot in a turbid medium follow a so- called modified Beer-Lambert law, the modifying term is the so-called Differential Path Length Factor ( DPF ) according to

J_ — „{-lia(A)dDPF(A,d)}

I₀ where I is the measured signal intensity at distance d from the source , I_o is the incident signal intensity at distance 0 , pa (X ) is the wavelength dependent absorption coefficient , and DPF ( X, d) is the correction factor, which depends on both wavelength and the lateral separation between illumination and collection point and accounts for the fact that the optical path length does not increase linearly with lateral separation in a heterogeneous medium . It has been observed that in a heterogeneous medium with absorption coefficient pa (X ) and the differential path length factor, the DPF itself is prone to modulation when the scattering characteristics of the heterogeneous medium are altered, e . g . , by increasing the blood Glucose concentration in human tissue . This process is reversible ; when the original scattering and absorption characteristics /coef f icients are reestablished, the DPF resumes its initial value ( s ) . The proposed principle facilitates the measurement of changes in the curvature with lateral distance of the (modified ) Lambert-Beer law, which is equivalent to measure the modulations of the DPL itself , during modulations of the heterogeneous medium by alteration of the scattering coefficient ( s ) / characteristics .

This aspect is given by the fact that in a heterogeneous medium consisting of a volume with refractive index and absorption coefficients nl ( X) and pal ( X) , respectively, interspersed with at least one additional medium of , say, refractive index n2 ( X) and -- in general -- absorption coefficient pa2 (X ) , also respectively, the propagation of light is not necessarily a straight path from the source- to detection point but undergoes a random walk due to scattering and reflection ( and absorption ) at the second medium contained within the first medium .

Accordingly, the path length travelled in medium 1 and/or medium 2 is no longer the straight line distance from the source- to detection point , but it turns into a statistical distribution of lengths according to a random walk, which leads to , on average , an increase of the distance travelled . The corresponding distance increase is described by the Differential Path Length Factor ( DPF) .

The inventor now observed that in a medium with changing scattering characteristics either as a function of z-height , as measured from the medium-to-ambient interface ( top surface ) , or as a function of separation between illumination and detection spot , or both, such as in human tissue and/or skin, the DPF is not constant . As a consequence , the modified Beer-Lambert law, when plotted over distance on a semi- logarithmic scale , is not a straight line but acquires a curvature , i . e . , a second-order- or quadratic term . This behavior is now used for the purpose of evaluating a substance within a liquid containing solid particles , whereas scattering occurs at the particles and the change of such scattering is given by the complex index difference between the refractive indices of the medium and the particles , respectively . Consequently, the curvature of signal strength with distance is a function of substance concentration . By evaluating the curvature of the signal strength with distance and its change over time , one can derive a concentration change of the substance .

However, the underlying measured signal still decays approximately with regard to the distance between emitter and detector . Hence , in particular at larger distances , the signal is difficult to measure due to its larger noise portion . On the other hand, simply increasing the light emission may not be sufficient , as detectors closer to the light source might saturate . Furthermore , as stated above , the curvature of signal strength with distance is a function of substance concentration . Consequently, it is preferable to measure nominal the same signal strength, such that any deviation is not or at least less caused by the exponential decay but by the substance concentration in the liquid .

The inventor now proposes a PPG measurement system comprising a light source and a detector, where the sensitive area of the detector scales exponentially with distance from the light source . It has been observed that the distance for skin type 2 and a normalized measured signal at 2 mm distance between light source and detector reduces to app 4% at 4 mm distance and further to . 4 % at 6mm. More particularly, the signal strength at 6mm has significantly decayed by a factor of approximately 233 . For a shot noise limited system, this transforms to a factor of ~15 ( sqrt ( 233 ) ) between the noise and signal portion .

The inventor therefore proposes a measured signal that is approximately constant at the at least three distances from the light source , such that the signal level and hence variability of the measurement in case of a shot-noise limited system is approximately the same for each distance and the assessment of curvature reduces effectively to evaluating a straight-line departure of the signal in between the near and the far site .

The inventor proposes in some aspects an innovative detector device for determining a substance concentration in a sample comprising liquid containing particles , in particular glucose in blood, wherein a refractive index of the liquid is dependent on a concentration of the substance dissolved therein and a density of particles in the liquid is substantially constant . The detector device comprises a housing with at least one light source and at least one detector or detecting component arranged therein . The housing comprises exits windows for accessing the at least one light source and at least one detecting component . Furthermore , the at least one detecting component within the housing is optically separated from the light emitting device .

The at least one light source is configured to emit light through an exit window onto a sample . Accordingly, the sample is usually placed above the exit window in the beam path of the light emitted by the at least one optoelectronic component . It is useful to place the sample tight onto the exit window to reduce any ambient light from reaching the surface of the sample . Consequently, in some aspects , the housing and/or the exit is configured to follow the shape of the sample or at least adj ust such that entry of ambient light is reduced .

Similar configurations may be applied to the entry window in front of the at least one detecting component . The entry window should be tight on the sample to avoid ambient light from getting through the entry window and reaching the detector component .

In accordance with the proposed principle , the detecting component comprises at least three detecting areas , each detecting area is arranged at different distances from the at least one light source .

The areas are configured to detect a light component corresponding to emitted light scattered through the sample . It is useful to place the sample tight onto the exit window to reduce any ambient light from reaching the surface of the sample . As an alternative , the detecting areas are located in positions resulting in detecting a light component corresponding to different path lengths of light travelling through the samples . In particular , the first path length is longer than the second path length, while the second path length is longer than a third path length, the third path length therefore being the shortest . For the purpose of this application, the expression path length is correlated to a distance on the detector component surface , from which light is emitted and a location at which light is detected .

As stated above , due to the decay of the emitted light at larger distances , a size of each detecting area is dependent on the wavelength of the emitted light and the distance to the at least one light source . The sizes of the respective areas are adj usted such that a signal-to- noise ratio of the detected light at each detecting area is in a predefined common interval .

A control circuit is coupled to the at least one light source and at least one detecting component , and particularly to the respective detecting areas thereof . It is configured to obtain at least three signals from said detecting component obtained from at least three detecting areas . In some aspects , the control circuit is configured to perform the steps not performed by the at least one light source and the at least one detecting component , namely the steps of deriving a -in particular maximal- deviation of the obtained signals or a processed curve thereof from a reference axis , said reference axis in particular including two of the at least three different signals . The control circuit is also configured to derive a substance concentration from derived deviation .

In some further aspects , the control circuit is configured to use the at least three obtained signals or processed curves thereof to derive a deviation from a reference line or value , said reference line or value in particular including two of the at least three different signals . In other words , the control circuit is configured to obtain a derivative extreme , e . g . , where the gradient of the measured values changes its sign . Finally, said control circuit is further configured to derive a substance concentration from the deviation .

Some aspects concern the predetermined common interval of the signal- to-noise ratio . It is generally proposed to maintain the signal-to- noise ratio at a comparable level by -as proposed, increasing the sizes of the respective detecting areas at increasing distances . Hence , it is proposed that the predefined common interval is around an average value of the at least three detected signals with its upper and lower limits less than 20% deviation from said average .

In some aspects , the sizes of the respective different areas may be different depending on the wavelength of the emitted light . For the green portion of the visible spectrum, a second detecting area being distanced to the light source by a factor of 2 in comparison to a first detecting area comprises a size that is between 20 and 30 times larger than the size of the first detecting area , and in particular 22 to 27 times larger . Likewise , a third detecting area is distanced to the light source by a factor of 3 in comparison to a first detecting area and comprises a size that is between 200 and 300 times larger than the size of the first detecting area , and in particular 220 to 265 times larger .

On the other hand, it has been found that the decay of light with increasing distance is smaller for light in the red and infrared portion of the spectrum . For instances , a second detecting area in such cases , which is distanced to the light source by a factor of 2 in comparison to a first detecting area comprises a size that is between 3 and 9 times larger than the size of the first detecting area , and in particular 3 . 75 to 7 times larger . Likewise , a third detecting area , which is distanced to the light source by a factor of 3 in comparison to a first detecting area , comprises a size that is between 12 and 25 times larger than the size of the first detecting area , and in particular 15 to 20 times larger .

In some aspects , it is suitable to provide for each detecting area comprising a plurality of sub-areas , said sub-areas having equal size and configured to be read out separately . This would allow to effectively increase or shrink the size of the detecting area at certain distances during readout . In this regard, the control circuit may be configured to address the respective sub-areas and read-out the respective signal , such that the signal-to-noise ratio for the detecting areas is substantially the same or at least within the common interval . Furthermore , some of the sub-areas can be used to detect the noise signal , when no light is emitted further improving the signal- to-noise ratio .

In some aspects , each detecting component comprises a plurality of area portions , said area portion substantially equally distributed around the at least one light source . For example , in some aspects , at least one of the detecting area is arranged along a virtual circle around the at least one light source . The detecting area may comprise subareas for example in the form of detector pixels that are arranged along a circle . These circles may comprise different radii corresponding to the detecting areas at different distances from the light source . In some other aspects , the at least one of the detecting area comprises at least two area portions , said area portions arranged on a virtual line through the at least one light source . These implementations ensure a symmetrical arrangement around the at least one light source , which may compensate for s kin irritations , surface contamination or other obstacles covering the detection area .

The detector device according to the proposed principle may comprise at least one light source that has a plurality of optoelectronic devices being arranged with different distances to the at least one detecting component ; and optionally comprising a ring shape or a quadratic shape optionally arranged, -particularly centrally- , around the at least one detecting component . With a quadratic shape , one integrates over a range of distances between "1" and "sqrt ( 2 ) This will most likely reduce the sensitivity of a curvature measurement , which leads to an increase in the variability of the measurement result . In some aspects , the control circuit is configured to control the at least one detecting component to detect a third light component while the at least one light source is not emitting . The third light component may include , for example , any ambient light , noise or other undesired component . The control circuit is configured to obtain least three signals utilizing the third light component and the light components corresponding to the different path . For instance , the third component can be subtracted from each of the at least three light components corresponding to emitted light scattered through the sample to remove ambient contribution including systematic errors . This will improve the signal-to-noise ratio .

In some further aspects , the at least one detecting component comprises a light filter comprising a low transmittance in a frequency spectrum different from a light spectrum emitted by the at least one light emitting component . This measure will further reduce ambient light portions in the detected component , thereby improving the quality of the detected component .

In some aspects , the control circuit is configured to read out subareas of the at least three detecting areas separately, such that the at least three detected signals from said sub-areas comprise a signal- to-noise ratio that varies from each with a deviation less than 20 % from an average .

The detector device can be implemented in a hand-held or a mobile device . It is possible to utilize already existing configurations , which are suitable for all kinds of PPG measurements , blood pressure or oxygen concentration, for example . In some aspects , a distance between the at least one light source and a first of the at least one detecting component is different to a distance between the at least one light source and a second of the at least one detecting component . Alternatively, a distance between a first of the at least one light source and the at least one detecting component may be different to a distance between a second of the at least one light source and the at least one detecting component . Both implementations will ensure that the optical path length the signal is travelling through the specimen is different , which will cause different amounts of scattering, thus leading to a difference in the measurable signal level per area . In some aspects , the inventor proposes a method for determining a substance concentration in a sample comprising liquid containing particles , in particular glucose in blood, wherein a refractive index of the liquid is dependent on a concentration of the substance dissolved therein and a density of particles in the liquid is substantially constant .

The method comprises in a first alternative , emitting, -in particular periodically- light of at least one wavelength onto a first location of the sample containing the liquid . The light can be a continuous light , but also provided in light pulses with known pulse length . This may be suitable to measure the ambient light portion during the off- period of the light pulse in order to subtract the ambient light portion from the cumulative light comprising scattered light from the onboard illumination plus the perturbing ambient light .

At least three different signals are obtained at respective second locations of the sample , each second location comprising a different distance to the first location and each obtained signal corresponding to a light component scattered by the sample . Consequently, this step proposes to conduct a plurality of measurements obtaining light components at various distances from the first location or more particularly from the incident light spot . As stated previously the signals obtained at the different locations contain some noise . Since noise is systematically added by the specimen ( living tissue ) , it is quite conceivable that the noise level at a larger distance is always greater than the noise level at a shorter distance . Hence , in other words , the proposed principle measures systematic noise at various locations . This noise is dependent on the distance of the signal travelled in tissue ( BOPL ) .

In accordance with the proposed principle , said signals are obtained by different detecting areas of a detecting component and said detecting areas comprise different sizes . The different sizes of the detection area cause a larger portion of light to be collected, such that a signal-to-noise ratio between the different detection areas is substantially equal . In some aspects , the overall signal portion of the signal collected at the different locations are substantially equal or at least similar for a subsequent processing .

The at least three obtained signals or a processed portion thereof are then used to derive an -in particular maximal- deviation from a reference line or values , said reference line or values in particular including two of the at least three different signals . The substance concentration is derived from the deviation .

The proposed principle results in a deviation from a reference directly correlated to the change of the substance concentration over time . This reference corresponds to a reference curvature of the measured signal across various distances . If no dependency over distance and substance concentration would be present , an actual measured curve should be similar to said reference curve . However , the curvature generated by the measured light components is not constant but deviates from the reference curve . Said deviation also changes with varying substance concentration in a specific manner . Hence , evaluation of the curvature generated by the measured light components , directly correlates to changes in the substance concentration . Due to the increased detector area at increasing differences , the signal-to-noise ratio or the signal strength is kept within a certain interval , thereby improving the overall measurement performance .

In some aspects , the sizes of the detecting areas are such that a signal-to-noise ratio of the at least three obtained signals deviates by less than 20% from an average value . At least two of the at least three different signals are obtained centrally around the first location .

It is suitable to further detect any ambient light and other effects to compensate for such signal portions later on . Consequently, a third light component during a duration, where no light is emitted, is detected at least at one of the second locations . The detection of the third component is set to a time frame , where no artificial light is emitted . The detected signals may be pre-processed, i . e . digitized using an AD converter , in particular having a converter having at least

14 bits of resolution .

In a similar manner, the light component obtained at the second location ( s ) is processed and may be amplified and converted to digital values . In some instances , the obtained light component from the second location ( s ) can be averaged prior to being processed to reduce smaller variations in the light component . For example , obtaining at least three different signals components can comprise obtaining a plurality of samples for each of the at least three different signals , and subsequently combining at least some of said samples .

Some further aspects concern the acquisition of data . For example , the first and/or second light pulse may comprise a periodicity between 20 Hz and 500 Hz , and in particular between 50 Hz and 200 Hz and particular above 75 Hz . Hence , an individual measurement is taken usually in a relatively short time span, in which a change of the substance concentration but also an external slower trigger ( e . g . a heartbeat ) does not affect the individual measurement . Furthermore , short pulses with high currents and correspondingly timed acquisition of the PD, increases the difference between signal generated by the onboard illumination and signal added by the ambient . In principle , this can help to reduce systematic error ( s ) and variability due to ambient light .

Each individual measurement is referred to as a sample , and several samples can be combined into a single value . However, several of such measurements may be influenced by the external parameter . If the frequency is high and the number of measurements taken is low, one can usually assume that a slow varying parameter , like the change of blood volume due to heartbeat , does not significantly affect the measurements . Otherwise , this can be taken into account by evaluating the effect of such slow varying parameter .

In addition, the first and/or second light pulse may comprise a duty cycle in the range of 1/20 to 1 /5 of the period in some aspects . This will reduce the overall power consumption, but is still sufficient to obtain the necessary information . As a result , the method can be implemented in a portable device , like a mobile phone , a watch or a portable medical device .

In some instances , the difference or the deviation is obtained by normalizing the measured values to a certain interval . In some aspects , the interval may be the range of [ 0 : 1 ] , wherein the limits are given by some measured values or processed values corresponding to one of the smallest distance . Consequently, in some instances , the at least three obtained signals are normalized such that measured signals for the smallest and the largest distance from the limits of the normalization interval .

Then, the deviation is determined based on the normalized values and more particularly from a virtual reference line through the limits . The maximum deviation found is strongly correlated to the substance concentration . Furthermore , the maximal deviation may be around half in between the limits , that is half the distance from the minimum distance to the maximum distance and vice versa . The deviation will vary with the concentration of the substance ; hence it is suggested in some instances to periodically repeat the measurements .

Some aspects concern the enhancement of the optical measurement by detecting ambient light and removing it from the actual measurement in order to improve the signal/noise ratio . In some instances , a third light component is detected during one or more time intervals , in which no light is emitted . This is referred to ambient light , and it is assumed that such ( constant ) light is present during emission of the light as well as when no light is emitted . The third light component is subtracted from the light component detected at the second location, which includes the actual scatter portion ( through the tissue ) as well as the ambient light portion .

SHORT DESCRIPTION OF THE DRAWINGS

Further aspects and embodiments in accordance with the proposed principle will become apparent in relation to the various embodiments and examples described in detail in connection with the accompanying drawings in which Figures 1A to 1C show some illustrations indicating the general model based on scattering of light due to difference in the refractive index as used in the present application;

Figures 2A to 2D illustrate several process steps of an embodiment for a method for determining a substance concentration in a sample comprising liquid containing particles in accordance with some aspects of the proposed principle ;

Figure 3 shows the results of the exemplary process of Figures 2A to 2D according to some aspects of the proposed principle ;

Figure 4A shows a reference curve of a glucose concentration in blood plasma at certain times after digesting sugar containing liquid;

Figure 4B illustrates several curves measured at certain times indicating in Figure 4A and specific distances in accordance with some aspects of the proposed principle ;

Figure 5 shows a time-measurement diagram with several measurements points as done in accordance with the proposed principle together with a corresponding reference curve ;

Figure 6 illustrates a time-measurement diagram with the same curves , but time shifted;

Figures 7A and 7B show various embodiments for a measurement device in accordance with some aspects of the proposed principle ;

Figures 8A to 8C show a top view of two more embodiments for a measurement device in accordance with some aspects of the proposed principle ;

Figure 9 illustrates yet another embodiment of a detector device in accordance with some aspects of the proposed principle . DETAILED DESCRIPTION

The following embodiments and examples disclose various aspects and their combinations according to the proposed principle . The embodiments and examples are not always to scale . Likewise , different elements can be displayed enlarged or reduced in size to emphasize individual aspects . It goes without saying that the individual aspects of the embodiments and examples shown in the figures can be combined with each other without further ado , without this contradicting the principle according to the invention . Some aspects show a regular structure or form . It should be noted that in practice slight differences and deviations from the ideal form may occur without , however , contradicting the inventive idea .

In addition, the individual figures and aspects are not necessarily shown in the correct size , nor do the proportions between individual elements have to be essentially correct . Some aspects are highlighted by showing them enlarged . However , terms such as "above" , "over" , "below" , "under" "larger" , "smaller" and the like are correctly represented with regard to the elements in the figures . So it is possible to deduce such relations between the elements based on the figures .

Figures 1A to 1C illustrate a model for the conceptual propagation of light in a liquid and heterogeneous medium . The medium comprises a refractive index of n=l as illustrated in Figure 1A . The incident light IL propagating through the liquid medium is scattered on a solid obj ect BC within the liquid . The obj ect can be a blood cell , for instance in the case of blood plasma as liquid medium. However , the model as well as the proposed principle explained later on is not limited to blood plasma , but also encompasses any other scattering obj ect having a refractive index of a larger than 1 in a liquid medium .

In the case of Figure 1A with n=l for the liquid medium and the scattering body having n>l , the incident light IL will partially be reflected at scattering body BS but also scatter thereupon with a large portion in the forward direction as indicated in the Figure . The amount of back scattering as well as forward scattering is based on the difference between the refractive indices in the liquid medium as well as on the scattering obj ect .

In Figure IB , the refractive index of the liquid medium is the same , that is n=l , while the scattering obj ect now comprises a different refractive index equal to the refractive index of the medium ( e . g . n=l as well ) . According to the conceptual propagation, this means that the incident light will not see any refractive index deviation between the scattering obj ect and the liquid medium . As a result , the incident light will propagate through the medium and not be scattered by the obj ect BC .

For Figure 1C , the liquid medium comprises a refractive index of larger to one , n>l . Here the refractive index of the scattering obj ect is equal to 1 , but more importantly, the refractive index of the scattering obj ect is smaller than the refractive index of the surrounding medium . Consequently, the situation is similar to the one depicted in Figure 1A, with the incident light being partially reflected backwards and partially scattered on the obj ect BC . The difference between the schedule back scattering and forward scattering is again based on the differences between the respective refractive indices .

In accordance with the model , the ratio of back scattering to forward scattering will change when the difference between the refractive indices ' changes . Such variation of the refractive indices and the difference between both indices occur when the concentration of a substance dissolved in the liquid medium changes . In the case of blood plasma , such substance is glucose , for example , but can also include other components or substances . The refractive index of blood plasma varies with varying concentration of glucose dissolved in it , and can consequently be used for evaluation of the overall absolute and/or relative glucose concentration .

For a measurement of the glucose concentration in the blood, it is assumed that the propagation of light from the light source to a light detector through tissue of a body and blood plasma will result in scattering and absorption of light , whereas the light takes a random path through the tissue . Depending on the scattering characteristics of the tissue , the path taken from the light source to the detector could be different . Particularly, scattering and absorption in the tissue are dependent on wavelength, with generally a higher absorption at shorter wavelength . Further , the amount of scattering is dependent on the distance between the light source and the detector when the resulting index contrast -that is , the difference of the refractive indices between the blood plasma as well as the red blood cells- is modulated depending on the glucose concentration .

It will change during consumption and digestion of glucose containing liquids or solids .

The inventor now proposes to utilize this effect to evaluate the scattered light components at different locations -that is , at different distances from the light source- over time . It has been observed that the scattering changes with a varying glucose concentration in blood at different distances from the light source . In other words , by evaluating a deviation between the signals at different distances over time , one can derive the glucose concentration therefrom.

However, it was observed that the measured signal decays with increasing distance between the emitter and the detector and detecting component . While this was expected, it should be noted that the overall System is limited with regard to shot noise . Shot noise is independent of the distance , such that it is substantially constant . Consequently, with increasing distance , the signal-to-noise ratio deteriorates . The inventor now proposes by to keep the signal-to-noise ratio substantially contact during a reference measurement at the various distances ( e . g . at the start of a longer glucose measurement ) . Alternatively, the inventor proposes to compensate for the signal decay by changing the size of the detecting component , such that the signal- to-noise ratio is substantially constant , or the measured signal is increased (by a factor equal to the signal decay over the distance ) . Since the signal-to— noise ratio for shot noise limited signals and signal strength correlate with each other, the term signal-to-noise ratio is used synonymous for both physical parameters throughout this application .

For the purpose of the following explanation as to the general method, it is assumed that the detector size has been adj usted depending on the wavelength and the respective distances between the emitter and the different detecting areas by a reference measurement CO ( at any given glucose concentration) . This measurement results in a straight curve with either no decrease at all ( corresponding to slope 0 ) or a given constant slope , as illustrated in Figure 2A. Figure 2B to 2 F show various process steps of the proposed principle to explain the basic principle in greater detail .

The respective Figures illustrate the distance on the X-axis between the light source and the detection point , ranging from 1 mm to about 8 mm . The Y-axis of the various Figures 2A to 2 F show different values for the obtained signals , including , but not limited to , signal levels and/or normalized values . For the purpose of explanation of the proposed principle , three different curves with various measurement points are obtained . The three different curves Cl , C2 and C3 represent optical measurements at different wavelengths , namely in the green, red and infrared portion . More particularly, curve Cl contains the measured values between the distance of measurements from a light source from 2 mm up to 7 mm obtained for light in the green spectrum . Curve C2 corresponds to the respective measurements for light in the red spectrum, curve C3 represents the measurement values for light in the infrared spectrum . Curve CO is a reference measurement curve for a given wavelength and with different detector sizes to ensure a constant signal decline over the various differences . As stated above , the normal decay is exponential . This curve allows compensating for such signal decay . As it can be seen, the curve varies slightly, but is substantially constant over the overall distance due to the increased detector size .

Figure 2A illustrates the three different curves Cl , C2 and C3 for three different wavelengths obtained at the distances between the source and the detecting components from 2 mm up to 7 mm . A finger with a light s kin tone ( Fitzgerald scale I I ) is placed on the sensor array, covering the light source and the detecting component . Light is emitted through the tissue and collected at various distances from the light source , namely at 2 , 3 , 4 , 5 , 6 and 7 mm . The raw signals are processed, i . e . the background noise has been subtracted to obtain the signal DC component shown in arbitrary units as illustrated in Figure 2A .

The curve Cl for the green light usually comprises the strongest decrease of almost overall three magnitudes of signal with increasing distance from 2 mm to 7 mm, However , due to the proposed principle of compensating the signal decay, the signal is almost constant over the various distances . The differences between the curves measured at different wavelengths are caused by different absorptions in the human tissue . In contrast to the measurement at green light , the curves C2 and C3 , corresponding to red and infrared light , respectively, do not show this strong absorption . They both show a similar pattern with only small fluctuation and deviations in between with a decrease from 2 mm distance up to 7 mm distance .

Still , all curves also experience a small gradient changing at approximately 4 mm distance between the light source and the detecting component . Measurement curve Cl with for green light comprises the strongest change in gradient , varying at around 4 mm, as visible from the Figure 2B . The same can be observed for the other curves C2 and C3 , although to a smaller extent . More particularly, the slope of all curves is different at distances from 2 mm to about 4 mm to 5 mm and then becomes larger from 5 mm to 7 mm . While the origin may be unknown, the gradient change is now evaluated and subsequently used to determine a change in the overall glucose concentration in accordance with the proposed principle . It has been observed that the gradient will change with varying glucose concentration .

To obtain the gradient , the following steps are exemplary performed as an exemplary embodiment for the method of determining substantive concentration in a sample comprising a liquid containing particles . Figure 2B shows the first process step, in which a normalization onto the largest value is performed . The values measured at the longest distances are normalized to a semi-logarithmic scale of 1 , thereby defining the first normalization points NP1 . Consequently, the points NP1 form the common first normalization point for the respective measured values and the corresponding curves Cl , C2 and C3 . They might be different now, although the deviation in between is probably exaggerated . In addition, the Y-axis is now in the correct scale , with the normalized last point NP1 equal to one . The Y-axis with the respective values has been normalized to a logarithmic scale with a value of 1 for the common first normalization point NP1 .

In the next step depicted in Figure 2C , the Y-axis is transformed using the logarithm for each of the curves separately such that all curves , provide a common normalization point NP1 , which is then set to 0 . While this step may reduce complexity in the processing effort , it can be omitted, that is the proposed method can be performed without normalization . The normalization process then continues with the next step depicted in Figure 2C . A second normalization with the measurement first value is performed, said value corresponding to the smallest distance between the light source and the detecting component of 2 mm . Those values for the three curves Cl , C2 and C3 are normalized to the value 1 . As a result , each curve now comprises a common first normalization point NP1 and a second normalization point NP2 , with the curves normalized to range between [ 0 : 1 ] . Hence , steps shown in Figures 2A to 2C correspond to a normalization of the received and pre-processed value to the interval of [ 0 : 1 ] , whereby the first and the last value correspond to the borders of the interval .

Figure 2E illustrates the results of the transformation normalizing all the measurement curves onto an interval of [ 0 : 1 ] with 0 being the first normalization point corresponding to the measured value at the largest distance and 1 corresponding to the measured value at the smallest distance between the light source and the detecting component . As it can be seen from the figure , there is a deviation from a straight line at around 4 mm . In a next step , a new reference curve RC is generated between the two normalization points . The reference curve is a virtual line during processing, it is shown herein for illustration purposes of the proposed method .

The reference curve is used to evaluate the deviation of the respective curves Cl , C2 and C3 and the corresponding values therefrom as depicted in Figure 3 . The data values exhibit a deviation from the reference curve obtained by subtraction of the respective measurements from the reference line . As shown in Figure 3 , there is a strong deviation at a distance of around 4 mm to about 5 mm present for all the three different wavelengths . The difference to the normalization points is about 6% to 8% of the normalized values . These process steps somewhat quantify the gradient change as already seen in the original measurements .

As observed, the maximum difference will vary over time when the glucose level in the blood plasma changes . More generally speaking, it has been found that the deviation of the signals and the curves derived therefrom from the reference curve defined by the normalized values will vary with changing concentration of a substance in the liquid, causing the refractive index of the liquid to change . Hence , the curvature of measured received light components across the plurality of distances becomes a function of substance concentration and/or a function of substance concentration change .

Such a relationship becomes most apparent at around 4 mm to 5 mm for blood plasma , but is visible also for larger or smaller distances . The reason for the variation at those specific distances for blood plasma may be caused by the tissue thickness and the distance from the surface of the tissue ( s kin surface ) to the blood containing area . It has further been observed that difference becomes the maximal after digesting sugar containing liquid or solid, resulting in a glucose change in the blood plasma of the consumer , but then decreases again as the glucose is consumed over time .

Figure 4A and 4B illustrate the results of an oral glucose tolerance test ( OGTT ) , in which a person consumes a sugar containing liquid at a specified time . Measurements are then taken at certain time intervals to evaluate the glucose level in the person' s blood plasma . In Figure 4A, a reference measurement has been taken, starting shortly after consumption at 14 . 00 . The reference measurement was performed by a reference glucose meter, using a finger-prick method followed by an electro-chemical sampling of a small blood portion from the person' s finger . As illustrated in the Figure , the glucose level rises sharply for about 30 minutes , remains at a maximal plateau for about 10 to 15 minutes and then slowly decreases to the initial value .

Figure 4B illustrates the respective measured curves for the optical distance scans at the respective times T1 to T5 also shown in the reference measurement in Figure 4A. The deviation from the straight line as previously obtained in accordance with the proposed principle strongly increases at the distance of 4 mm to 5 mm, showing the maximum deviation of difference with an increasing glucose level in the blood plasma . During the given time frame between T1 and T5 , the deviation approximately doubles . Consequently, the results of the processing method, namely the deviation from the straight line during the initial increase of glucose concentration following the oral glucose tolerance test indicate the correlation to the incomplete increasing blood sugar concentration .

Figure 5 and Figure 6 illustrate curves as shown in the previous Figure 4 for a longer time after digesting a sugar containing liquid . The figures illustrate measured reference curves RC taken by the above- mentioned glucose reference meter , as well as a quadratic term of a polynomial fit of optical scan measurements for a wavelength of 940nm . The X-axis illustrates a time axis for the entire duration of the oral glucose tolerance test . The wavelength of 940 nm provides a better signal-to-noise ratio compared to the green or red light , respectively .

Figure 5 illustrates the nominal timescale for both reference and optical measurement . It is observed during a variety of measurements that the optical measurement and the evaluation of the corresponding glucose level will show an increase of the glucose level ( or at least a change in the refractive index of the blood plasma ) while the reference measurement indicates the same glucose increase with a delay of appr . 10 minutes to 20 minutes . In the present exemplary measurement of Figure 5 , the time delay of the glucose increase between the optical measurement and the reference measurement is approximately 19 minutes . Each optical measurement is taken at 4 mm distance , where the difference of the normalized measurements from the reference curve illustrated in Figures 2A to 2D is maximal . While the optical measurement is somewhat noisy due to the experimental setup, the correlation between the reference measurement and the optical measurement is clearly visible .

Figure 6 illustrates the same two curves after bringing the reference measurement forwarded by 19 minutes . As illustrated, the optical measurement and the reference measurement are strongly correlated, apart from some fluctuations and noise being present in the optical measurement data . The results suggest that the non-invasive optical measurement responds sooner to changes of glucose concentration in the blood plasma compared to the reference device and the finger prickmethod .

Figure 9 shows a detector device for optical measurement of a substance concentration in a liquid medium, wherein the liquid medium contains several solid particles . The detector device can be implemented in existing handheld devices , like for example mobiles , tablets , but also directly assembled as a medical diagnostic device . The detector device comprises a fingerprint sensor 1 , or finger clip covering the tip of the finger . As an alternative , the sensor 1 can be implemented as a button to be glued over a part of the s kin, like on an arm, ear and the like . In some instances , the sensor can be an ear clip , smart watch, fitness tracker, finger ring and the like . The sensor is powered from the control and supply unit 2 , which can be battery driven, but also coupled to the main power grid . Sensor 1 is connected to power unit 2 via an interface , receiving power and control signals therefrom . The control unit 2 receives the measured signals and calculates the glucose level in accordance with the proposed principle using the deviation from a reference curve , or the extreme value of the first derivative of the measured values . Finally, the results can be saved or displayed to a user for review . For this purpose , the control unit is connected to terminal and storage device 3 .

Figure 7A and 7B illustrate two embodiments for a sensor configured to perform the method in accordance with the proposed principle .

Sensor 1 comprises a substrate 30 on or in which a plurality of detecting components 20 , 21 and 22 are implemented . Although only three detecting components are shown herein, a plurality of detecting components can be used, as indicated by the dots in the drawing . The detectors can be arranged in an array of rows and columns , but also comprise a different structure as outlined further below . The detecting components 20 , 21 and 22 are distanced from a light source 10 by well- defined distances . Depending on the number of detecting components used, a plurality of measurements at different distances can be obtained . However , at least three sites or distances are required to evaluate the curvature of the signal decay with distance .

An optical barrier 40 separates the detecting components 20 , 21 and 22 from light source 10 being implemented in sensor 1 . Light source 10 is configured as a light emitting diode or more generally as an optoelectronic device . The light source can emit light of a single wavelength or a plurality of different wavelengths . The expression "single wavelength" corresponds to emitted light with a relatively small band having an FWHM of appr 30 nm to 70 nm. In some aspects , the light source can be a laser emitting light with a very small bandwidth . Light emission of a narrowband is beneficial , if the detecting components are implemented with narrow filters blocking light of different wavelengths . This will reduce ambient light and improve signal/noise ratio .

The principal realization of a detector device suitable for such measurement , which for example , can be implemented in a display device , is illustrated in Figure 8A. The detector device is integrated in a display device , which comprises a plurality of pixels arranged in row and columns . The display and pixels sizes are not subj ect to scale , but it shows the general principle . Each pixel in this embodiment contains a detector as well as emitting pixels of different colors . The detectors can be read out separately and then combined, such that a plurality of detectors forms a detecting area or sub-regions thereof .

A central Pixel Pl ( or a plurality therefore ) is emitting green light or any other light of a certain color . As illustrated, several detectors around the emitting pixels are activated . Those detectors are selected such, that they lie in a virtual circle around the centrally arrangement emitting pixel at various distances .

The activated detectors along each virtual circle belong to the same detecting area . More particularly, the activated detectors are substantially positioned symmetrically long said virtual circle in order to compensate for possible signal fluctuations along a certain direction . Each photodetector has a specific size and provides a specific photo current , although the level of the detected photo current is dependent on the distance to the emitter Pl .

As shown in Figure 8A, the number of activated photo detectors as subareas of the detecting area along the virtual circles increase with increasing distances . the number of activated photodetectors along each circle can be adj usted and varied by the control unit , such the sum of the detected photo current and/or the signal-to-noise ration of the detected photo current in each circle is substantially equal . In other words , the control circuit conducts a reference measurement , or utilizes a predetermined pattern of detectors . During such reference measurement , detectors along the virtual line are activated or deactivated such that the photo current at a given distance ( i . e . along the respective virtual lines ) are substantially equal and/or the signal-to-noise ratios at said given distances are substantially equal . Afterwards , the measurement can start and detect a change in the substance concentration using the detected signal at various differences as explained above . In the embodiment of Figure 8A, a plurality of measurements at various are evaluated, improving the overall accuracy . The reference measurement also allows using different lights , as detectors in each of the pixels can be read out separately . As the signal decay is also wavelength dependent , such implementation can perform measurements in various portions of the visible spectrum .

Figure 8A illustrates an embodiment of a display using the detectors along a virtual circle . However, this is not necessary, and one can utilize different detector shapes , particularly when the generated error is small . Figure 8B illustrates another exemplary embodiment , in which a light emitter Pl capable of emitting light of different colors is arranged centrally around a three different detector areas DI , D2 and D3 . The detector arrays are arranged as rectangles around the emitter Pl . This will result in some slight deviation along the edges of each rectangle , as the distance to an edge is longer than the distance to the line center of each rectangle . However, this can either be compensated based on the sensitivity of the detector or by the fact that it occurs for each rectangle . As an alternative , hexagonal patterns of detecting areas can be used . The area of the different varies and increases with increasing distances . For example , for red light emitted by pixels Pl , the detector size of detector D2 is appr . 5 times the size of detector DI with the size of detector D3 being another 3 to 4 times larger than the size of detector D2 .

Figure 8C illustrates a further embodiment of a detector arrangement in accordance with the proposed principle , suitable to perform the measurement for green light . In this embodiment the detector arrays are arranged in a virtual cross with the center being the emitter Pl . Each detecting area comprises 4 sub-regions arranged at 2mm, 4mm and 6mm, respectively along the virtual cross line . As shown, the area sizes between the respective sub-regions and the detecting areas differ by a factor of 25 and 233 , respectively, allegedly following the exponential signal decay . The first detecting component DI comprises a size of 0 . 2 x 0 . 2 mm2 at distance 2 mm from the emitter Pl . At four millimeters detector sub-regions with an active area of 1 . 25 x 0 . 4 mm2 , which provides an increase in active area of 25x, and similarly four detectors D3 with 2 . 3 xl mm2 , are placed at 6 mm distance , which will have a 230x larger area than the first Detector DI . Referring back to Figures 7 and 8 in operation of the sensor 1 , the light source 10 will emit light of one or more wavelengths , for example , infrared light or red light , which is transmitted into the skin of the finger . The light is either partially absorbed or scattered randomly within the tissue . Some of the scattered light components will reach the first detecting component 20 , while other light components are scattered in the tissue reaching the detecting components 21 and 22 , respectively . Due to the increasing distance , the light component received by the respective detectors comprise different intensities due to the additional absorption . The detecting components 20 , 21 and 22 are connected to amplifier units and analogue-to-digital converters to process the received light components scattered within the tissue .

The respective received signals are evaluated and processed in accordance with the method disclosed herein .

Figure 7B shows a slightly alternative embodiment compared to the embodiment of Figure 7A with a plurality of light sources 10 , 11 and 12 being arranged at various distances to a common detecting component 20 . Similar to the previous embodiment , the number of light sources can be varied as well as the location with regard to the detector . Further, the distance between the detecting component 20 and the optoelectronic device 10 closest to the photodetector is approximately 2 mm distance . An optical barrier 40 is arranged between the first light source 10 and the detecting component 20 .

In the embodiment , the various light sources 10 , 11 , and 12 will emit light through the tissue of the finger (wrist , ear lobe , in-ear, chest or other easy accessible body parts with blood vessels close by) including blood vessels therein scattering in various ways , whereas a portion of the light will reach the respective detecting component 20 . In contrast to the previous solution, however, the light sources 10 , 11 and 12 emit light at different times ( e . g . periodically and subsequent ) to be able to distinguish between the different light sources and therefore between the different distances , on which scattered light is received by the photodetector 20 . In an alternative solution, the light emitted by the light sources 10 11 and 12 is modulated in different ways . The detector will receive the scattered light portion of all light sources simultaneously . By proper demodulation and/or signal filtering , the different received light components are separated and then further processed, i . e . converted into digital values . Similar to the previous solution, the photodetector will amplify the received signal . Signal separation into the different components can be performed prior to digital conversion or afterwards , that is , the amplified measured combined signal is first converted into a digital signal and subsequently separated into the three different components .

The important aspect in the design of a curvature detector is the fact that signals decay nearly exponentially according to Lambert-Beer' s law, where the argument of the exponential function depends primarily on wavelength and skin type . Thus , one can either produce one specific detector arrangement that is accurate for one wavelength and one s kin type , or choose a compromise that works well enough for several wavelengths and/or s kin types . The embodiment with separately small detectors that can be combined for read-out are a suitable implementation .

In each of those cases , it is ensured that the detecting area closest to the light source does not saturate while the detecting area farthest away still receives enough light , so that signal variability is not a limiting factor . Similar outcomes with a same sized detecting area can also be realized by making sequential measurements , where the signal at 2mm is measured, e . g . , by turning on a single LED once , or for one period of time , at 4 mm turning on the emitter 26 times , or for a period that is 26 times as long as the first period, and having an acquisition system that is capable of integrating the generated photocurrent , and 233 times for the detecting array at 6 mm. This principle can be generalized, simply that the emitter is turned on for a long period of time and the detecting areas only read the photo-current during certain time portions thereof . However , while relatively easy to implement , this approach means that the measurements are no longer strictly performed at the same time , and the power consumption increases by a factor of 259 over the single measurement .

In the same way, it is also possible to increase the current through the emitter to generate more light when measuring at longer distances .

However, there are several challenges , including that the output of the emitter may be limited due to thermal overrun, or that a larger emitter is required to keep the current density low enough . Further, a driver with higher current capacity may be needed and with a more dynamic range . Nevertheless , the proposed approach is not limited to a single skin type or three different distances . Rather, one can design a curvature detector with suitable choices for ratios of the active areas , ratios of the number of pulses , ratios of currents , and/or any combination thereof , that satisfies the needs of the application .

LIST OF REFERENCES

1 sensor 10, 11, 12 light source

20, 21, 22 photo detector

30 substrate

40 optical barrier

Cl, C2, C3 measurements signals Tl, T2, T3 measurement times

T4, T5 measurement times

BC scattering body, blood cell

NP1, NP2 normalized point

Claims

1 . Detector device for determining a substance concentration in a sample comprising a liquid containing particles , in particular glucose in blood, wherein a refractive index of the liquid is dependent on a concentration of the substance dissolved therein and a density of particles in the liquid is substantially constant , said detector device comprising :

An arrangement having at least one light source and at least one detecting component , wherein the at least one detecting component within the housing is optically separated from the at least one light source ; wherein said at least one light source is configured to emit light through an exit window onto a sample ; and wherein said detecting component comprises at least three detecting areas , each detecting area arranged at different distances from the at least one light source ; wherein each area is configured to detect a light component corresponding to emitted light scattered through the sample ; wherein a size of each detecting area is dependent on the wavelength of the emitted light and the distance to the at least one light source , such that a signal level and/or a signal-to- noise ratio of the detected light at each detecting area is in a predefined common interval ; a control circuit coupled to the at least one light source and at least one detecting component and configured to obtain at least three signals from said detecting component obtained from at least three detecting areas ; said control circuit further configured to derive a -in particular maximal- deviation of the at least three obtained signals or a processed curve thereof from a reference axis , said reference axis in particular including two of the at least three different signals ; said control circuit further configured to derive a substance concentration from derived deviation .

2 . Detector device according to claim 1 , wherein the predefined common interval is approximately around an average value of the at least three detected signals , optionally with its respective limits deviating less than 20% of said average value .

3 . Detector device according to any of the preceding claims , wherein for an emitting light in the green portion of the visible spectrum; a second detecting area being distanced to the light source by a factor of 2 in comparison to a first detecting area comprises a size that is between 20 and 30 times larger than the size of the first detecting area, and in particular 22 to 27 time larger ; and/ or a third detecting area being distanced to the light source by a factor of 3 in comparison to a first detecting area comprises a size that is between 200 and 300 times larger than the size of the first detecting area, and in particular 220 to 265 time larger .

4 . Detector device according to any of the preceding claims , wherein for an emitting light in the red or infrared portion of the visible spectrum; a second detecting area being distanced to the light source by a factor of 2 in comparison to a first detecting area comprises a size that is between 3 and 9 times larger than the size of the first detecting area , and in particular 3 . 75 to 7 time larger ; and/ or a third detecting area being distanced to the light source by a factor of 3 in comparison to a first detecting area comprises a size that is between 12 and 25 times larger than the size of the first detecting area, and in particular 15 to 20 time larger .

5 . Detector device according to any of the preceding claims , wherein each detecting area comprises a plurality of sub-areas , said subareas having equal size and configured to be read out separately .

6 . Detector device according to any of the preceding claims , wherein each detecting area comprises a plurality of area portions , said area portion substantially equally distributed around the at least one light source .

7 . Detector device according to any of the preceding claims , wherein

- at least one of the detecting areas is arranged along a virtual circle around the at least one light source ; or

- at least one of the detecting area comprises at least two area portions , said area portions arranged on a virtual line through the at least one light source .

8 . Detector device according to any of the preceding claims , wherein said at least one light source comprises a plurality of optoelectronic devices being arranged with different distances to the at least one detecting component ; and optionally comprising a ring shape or a quadratic shape optionally arranged, -particularly centrally- , around the at least one detecting component .

9 . Detector device according to any of the preceding claims , wherein, the at least one detecting component is configured to detect a third light component while the at least one light source is not emitting .

10 . Detector device according to any of the preceding claims , wherein the at least one detecting component comprises a light filter comprising a low transmittance in a frequency spectrum different from a light spectrum emitted by the at least one light source .

11 . Detector device according to claim 1 , wherein the control circuit is configured to read out sub-areas of the at least three detecting areas separately, such that the at least three detected signals from said sub-areas comprise a signal-to- noise ratio that varies from each with a deviation less than 20 % from an average ; or the control circuit is configured to scale a size of the subareas to be read-out or a size of the at least three detecting areas depending on at least one of wavelength and s kin type .

12 . A method for determining a substance concentration in a sample comprising a liquid containing particles , in particular glucose in blood, wherein a refractive index of the liquid is dependent on a concentration of the substance dissolved therein and a density of particles in the liquid is substantially constant , comprising the steps of :

- emitting , -in particular periodically- light of at least one wavelength onto a first location of the sample containing the liquid;

- obtaining at least three different signals at a respective second location of the sample , each second location comprising a different distance to the first location, each obtained signal corresponding to a light component scattered by the sample ; wherein said signals are obtained by different detecting areas of a detecting component and said detecting areas comprise different sizes ;

- deriving a -in particular maximal- deviation of the obtained signals or a processed curve thereof from a reference axis , said reference axis in particular including two of the at least three different signals ;

- deriving a substance concentration from the deviation .

13 . Method according to claim 12 , wherein the sizes of the detecting areas are such that a signal-to-noise ratio of the at least three obtained signal deviates by less than 20% .

14 . Method according to claim 12 or 13 , wherein at least two of the at least three different signals are obtained centrally around the first location .

15 . Method according to any of the preceding claims , wherein the first and/or second light pulse comprises a periodicity between 20 Hz and 500 Hz , and in particular between 50 Hz and 200 Hz and particular above 75 Hz ; and wherein, optionally, the first and/or second light pulse comprises a duty cycle in the range of 1/20 to 1/5 of the period .

16 . Method according to any of the preceding claims , wherein obtaining at least three different signals component comprises obtaining a plurality of samples for each of the three different signals , and subsequently combining at least some of said samples .

17 . Method according to any of the preceding claims , deriving a -in particular maximal- deviation of the obtained signals comprises : normalizing the at least three obtained signals , in particular , with an obtained signal corresponding to the smallest distance and an obtained signal corresponding to the smallest distance being the limits of the normalization interval ; obtaining the maximum deviation from a virtual reference through the limits of the normalization .

18 . Method according to any of the preceding claims , further detecting a third light component during one or more time intervals , in which no light is emitted; and wherein optionally the step of obtaining at least three signals comprises subtracting the third light component from light components detected at the second location .