Non-invasive monitoring method of fluorescent tracer with background separation correction

Abstract

A method for non-invasively monitoring fluorescent tracers with background separation correction is disclosed, particularly a method for monitoring time-varying fluorescence emitted from a fluorescent agent within a diffusely reflecting medium having time-varying optical properties. The method comprises providing at least two measurements obtained from a patient before and after administration of the fluorescent agent, the measurements comprising a Flr adjacent to the medium detected by a filtered light detector during illumination of the medium with light of an excitation wavelength. meas signal, and at least one DR signal, the DR signal is selected from: #imgabs0#DR em and DR em,filtered The method further includes identifying a post-balanced portion of the measurement data set and assigning each Flr within the post-balanced portion of the measurement data set meas Signal conversion to IF agent signal, the IF agent The signal represents the intensity of the detected fluorescence emitted only by the fluorescent agent in the medium. The disclosed method includes eliminating the effects of leakage of excitation level light and meas Eliminate the influence of autofluorescence in the signal.

Description

[0001] This application is a divisional application of the Chinese national phase application of the PCT application with a filing date of January 30, 2018, an international application number of PCT / US2018 / 016053, and an invention name of “Non-invasive monitoring method of fluorescent tracers with background separation correction”, which entered the Chinese national phase on July 29, 2019 and has an application number of 201880009056.X.

[0002] Cross-reference to related applications

[0003] This application claims the benefit of U.S. Provisional Application No. 62 / 452,021, filed January 30, 2017, which is incorporated herein in its entirety. Technical Field

[0004] The present disclosure generally relates to methods for non-invasively monitoring fluorescent tracers in a medium characterized by light scattering and / or absorption. More specifically, the present disclosure relates to methods for non-invasively assessing renal function by monitoring the clearance of exogenous fluorescent tracers from tissues in a patient. Background Art

[0005] In order to minimize the risk of acute renal failure caused by various clinical, physiological and pathological conditions, it is very necessary to carry out real-time dynamic monitoring of the patient's renal function at the bedside. It is particularly important in the case of critically ill patients or injured patients because a large proportion of these patients are faced with the risk of multiple organ failure (MOF) caused by one or more serious functional disorders, such as: acute lung injury (ALI), adult respiratory distress syndrome (ARDS), high metabolism, hypotension, persistent inflammation and / or sepsis. Renal function may also be impaired due to the renal damage associated with the nephrotoxic drugs used as a part of surgery (such as angiography, diabetes, autoimmune diseases and other functional disorders and / or damage related to renal damage). In order to assess the status of the patient and monitor the severity and / or progression of renal function for a long time, there is considerable interest in developing a simple, accurate and continuous method to determine renal failure, preferably by non-invasive procedures.

[0006] Serum creatinine concentration, an endogenous marker of renal function, is typically measured from a blood sample and used in conjunction with patient demographic factors such as weight, age, and / or race to estimate glomerular filtration rate (GFR), a measure of renal function. However, creatinine-based renal function assessments can be prone to inaccuracies due to a number of potential factors, including age, hydration status, renal perfusion, muscle mass, dietary intake, and many other anthropometric and clinical variables. To compensate for these differences, a series of creatinine-based equations (recently expanded to cystatin C) have been developed that include factors such as sex, race, and other relevant factors for estimating glomerular filtration rate (eGFR) based on serum creatinine measurements. However, these eGFR equations do not provide any means of compensating for most of the above-mentioned sources of variation and therefore have relatively poor accuracy. Furthermore, eGFR methods typically produce results that lag 72 hours behind the true GFR.

[0007] Existing methods for measuring GFR have used substances such as inulin, phthalates, 51 Cr-EDTA, Gd-DTPA and 99m Other endogenous labels, such as o-iodine hippurate or 99m Tc-MAG3 123 I and 125 I, has been used in other existing methods for assessing tubular secretory processes. However, the use of typical exogenous labeled compounds can be accompanied by various undesirable effects, including the introduction of radioactive materials and / or ionizing radiation into the patient's body, as well as laborious ex vivo processing of blood and urine samples, making existing methods using these exogenous markers unsuitable for real-time monitoring of renal function at the patient's bedside.

[0008] The availability of real-time, accurate, and reproducible measurements of renal excretion rates using exogenous markers in patient-specific, potentially variable settings would represent a significant improvement over any currently practiced method. Furthermore, methods that rely solely on renal clearance of exogenous chemical entities would provide direct and continuous pharmacokinetic measurements without the need for subjective interpretation based on age, muscle mass, blood pressure, etc. Summary of the Invention

[0009] The present application provides a method for monitoring time-varying fluorescence emitted from a fluorescent agent in a diffuse reflective medium having time-varying optical properties, the method comprising: providing, by a monitoring system, a measurement data set comprising a plurality of measurement data entries, each measurement data entry comprising at least two measurement values obtained from a patient at a data acquisition time before and after administration of an exogenous fluorescent agent, the at least two measurement values comprising a fluorescence emission signal Flr detected by a filtered light detector at a third region adjacent to the diffuse reflective medium during illumination of the diffuse reflective medium by excitation wavelength light from a first region, and at least one diffuse reflectance signal DR, the at least one diffuse reflectance signal DR being selected from: DR ex A signal DR is detected by an unfiltered light detector at a second region adjacent to the diffusely reflective medium during illumination of the diffusely reflective medium by the excitation wavelength light from the first region adjacent to the diffusely reflective medium; em signal, detected by the unfiltered light detector at the second region during the period when the emission wavelength light from the first region illuminates the diffuse reflective medium; and DR em,filtered signal, detected by the filtered light detector at the third area during the period when the emission wavelength light from the first area illuminates the diffuse reflective medium; and identifying the post-reagent application portion of the measurement data set by the monitoring system; converting the fluorescence emission signal Flr signal of each measurement data entry within the post-reagent application portion of the measurement data set into a corrected fluorescence signal by the monitoring system, wherein the conversion includes eliminating the effect of leakage through of the excitation wavelength light into the fluorescence emission signal Flr and eliminating at least one of the effect of the autofluorescence contribution from the fluorescence emission signal Flr.

[0010] The present application also provides a method for determining renal function of a patient, the method comprising: providing, by a monitoring system, a measurement data set comprising a plurality of measurement entries, each measurement data entry comprising at least two measurement values obtained from tissue of the patient at corresponding data acquisition times before and after administration of an exogenous fluorescent agent, the at least two measurement values comprising a fluorescence emission signal Flr detected by a filtered light detector at a third region adjacent to a diffuse reflective medium during illumination of the diffuse reflective medium by excitation wavelength light from a first region, and at least one diffuse reflectance signal DR, the at least one diffuse reflectance signal DR being selected from: DR ex A signal DR is detected by an unfiltered light detector at a second region adjacent to the diffusely reflective medium during illumination of the diffusely reflective medium by the excitation wavelength light from the first region adjacent to the diffusely reflective medium; em signal, detected by the unfiltered light detector at the second region during the period when the emission wavelength light from the first region illuminates the diffuse reflective medium; and DR em,filtereda fluorescence emission signal Flr detected by the filtered light detector at the third region during the period when the emission wavelength light from the first region illuminates the diffuse reflective medium; and identifying a post-reagent administration portion of the measurement data set; and converting the fluorescence emission signal Flr of each measurement data entry within the post-reagent administration portion of the measurement data set into an intrinsic fluorescence signal IF by a monitoring system. agent , the IF agent represents the corrected fluorescence intensity emitted by the exogenous fluorescent agent from the diffuse reflective medium; the intrinsic fluorescence signal IF agent Fitting the rate of change during different data acquisition times to automatically identify the post-equilibrium portion of the measurement data set; using the intrinsic fluorescence signal IF from the post-equilibrium portion of the measurement data set through a monitoring system agent The rate of change determines renal function. BRIEF DESCRIPTION OF THE DRAWINGS

[0011] The present disclosure will be better understood when considering the following detailed description thereof, and features, aspects, and advantages in addition to those set forth above will become apparent. This detailed description refers to the following drawings, in which:

[0012] Figure 1 is a schematic diagram of a single wavelength renal monitoring device in one aspect;

[0013] Figure 2 is a schematic diagram of a dual wavelength renal monitoring system in one aspect;

[0014] Figure 3 is a graph summarizing the absorption, transmission, and emission spectra of various devices, materials, and compounds relevant to the non-invasive monitoring of exogenous fluorescent agents in vivo within a defined range of optical wavelengths from about 430 nm to about 650 nm;

[0015] Figure 4 is a graph summarizing the absorption spectra of oxyhemoglobin (HbO2) and deoxyhemoglobin (Hb) defined within the wavelength range of light from about 200 nm to about 650 nm;

[0016] Figure 5 is a schematic diagram of the timing of light pulse cycles associated with data acquisition in a dual wavelength renal monitoring system in one aspect, wherein each light pulse cycle sequentially includes light pulses generated at an excitation wavelength and an emission wavelength;

[0017] Figure 6 is a side view of a sensor head of a renal function monitoring system in one aspect;

[0018] Figure 7 yes Figure 6 Bottom view of the sensor head.

[0019] Figure 8 yes Figure 6 A top view of the interior of the sensor head. Figure 8 shows the arrangement of various electrical components within the housing of a sensor head of a renal function monitoring system in one aspect;

[0020] Figure 9 yes Figure 8 An enlarged view of the interior.

[0021] Figure 10 is a schematic diagram of a hole formed in a contact surface of a sensor head of a renal function monitoring system in one aspect;

[0022] Figure 11 is a schematic diagram of synchronous detection of light by light detectors of a sensor head in one aspect;

[0023] Figure 12 is a schematic diagram of optical signal modulation and demodulation performed by a sensor head in one aspect;

[0024] Figure 13 is a block diagram illustrating subunits of a processing unit in one aspect;

[0025] Figure 14 is a plot of the raw fluorescence signal as a function of time, showing the various phenomena that contribute to the total signal;

[0026] Figure 15 is a graph of the intrinsic fluorescence signal as a function of time with and without autofluorescence correction, illustrating the effect of autofluorescence correction on the renal decay time constant (RDTC) derived from the analysis of the intrinsic fluorescence signal;

[0027] Figure 16 is a graph of the raw fluorescence signal as a function of time, where the final fluorescence signal is lower than the raw background fluorescence signal level due to various phenomena that contribute to the total signal;

[0028] Figure 17A is a graph of the raw fluorescence signal and excitation light leakage as a function of time;

[0029] Figure 17B yes Figure 17A The graph of the original fluorescence signal and the corrected fluorescence signal, the corrected fluorescence signal including the elimination of Figure 17A The original fluorescence signal of the excitation light leakage;

[0030] Figure 18 is a graph comparing the original fluorescence signal (blue line) and the autofluorescence signal (green line) obtained before the injection of exogenous fluorescent agent;

[0031] Figure 19A Comparison of the original fluorescence signal obtained before injection of exogenous fluorescent agent Autofluorescence signal DR em and diffuse reflection signal DR em,filtered Graph of

[0032] Figure 19B Comparison of the original fluorescence signal obtained after injection of exogenous fluorescent agent Autofluorescence signal DR em and diffuse reflection signal DR em,filtered Graph of

[0033] Figure 20 is a flow chart outlining the steps of a background correction method for removing the effects of excitation wavelength light bleed-through and autofluorescence from a raw measured fluorescence signal;

[0034] Figure 21 Representative raw fluorescence signal measurements (IF) detected by the renal monitoring device obtained before and after injection of exogenous fluorescent agent agent ) curve graph;

[0035] Figure 22A is a block diagram illustrating a plurality of modules of a pre-processing subunit in one aspect;

[0036] Figure 22B is a block diagram illustrating a plurality of modules of a pre-processing subunit in the second aspect;

[0037] Figure 23 is an isometric view of a sensor head of a renal function monitoring system in a second aspect;

[0038] Figure 24 yes Figure 23 A bottom view of the sensor head of the renal function monitoring system is shown.

[0039] Figure 25 yes Figure 23 An isometric view of the sensor head of the renal function monitoring system is shown with the upper housing and various electrical components removed to expose the inner housing;

[0040] Figure 26 yes Figure 25 Exploded view of the inner housing of the sensor head shown.

[0041] Figure 27 It is shown that without the application of exogenous fluorescent agents for more than a day and Flr meas Graph of

[0042] Figure 28 Figure 2 shows the images immediately before and after application of exogenous fluorescent agent. and Flrmeas a graph of ; and

[0043] Figure 29 is the Flr determined by general experience leakt hroug h and from a database of 33 patients A graph showing the relationship between .

[0044] This written description uses examples to disclose the invention, including the best mode, and also to enable those skilled in the art to practice the invention, including making and using any equipment or systems and performing any combined methods. The patentable scope of the invention is defined by the claims and may include other examples that occur to those skilled in the art. If these other examples have structural elements that do not differ from the literal language of the claims, or if they include equivalent structural elements that do not differ substantially from the literal language of the claims, then these other examples are intended to fall within the scope of the claims. DETAILED DESCRIPTION

[0045] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this disclosure belongs. Although any methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present disclosure, preferred materials and methods are described below.

[0046] As used herein, a sample refers to a single discrete data value acquired from a signal and / or telemetry analog-to-digital converter (ADC) for a single acquisition / telemetry channel.

[0047] As used herein, a measurement refers to a single discrete data value created by demodulating or accumulating a sequence of samples from one acquisition channel.

[0048] As used herein, a measurement value refers to a collection that includes demodulated in-phase, demodulated out-of-phase, and averaged measurement values from one acquisition channel.

[0049] As used herein, a measurement value subset refers to a set including all measurement values of all acquisition channels during illumination of a single source LED. For example, all measurements of an acquisition channel may include demodulated in-phase, demodulated out-of-phase, and average measurements.

[0050] As used herein, a measurement set refers to a set that includes one measurement subset for each source LED.

[0051] As used herein, acquisition refers to the overall process of obtaining a set of measurements.

[0052] As used herein, a measurement sequence refers to a sequence of one or more measurement sets.

[0053] As used herein, a telemetry value refers to a single discrete data value acquired from a single channel of a telemetry ADC.

[0054] As used herein, a telemetry set refers to a set that includes one telemetry value from each telemetry channel.

[0055] As used herein, a diffusely reflective medium refers to any material through which light propagates that includes a plurality of moieties, particles, or molecules that scatter, reflect, and / or absorb light as it propagates. The distribution of the moieties, particles, and / or molecules can be uniform or non-uniform and can vary over time.

[0056] In various aspects, disclosed below are systems and methods for monitoring time-varying fluorescence emitted from a fluorescent agent within a diffusely reflective medium having time-varying optical properties. In one aspect, disclosed are systems and methods for monitoring time-varying fluorescence emitted from an exogenous fluorescent agent within patient tissue. The systems and methods in this regard may be used in a variety of contexts, including, but not limited to, monitoring renal function in a patient in real time by monitoring the reduced fluorescence emitted by an exogenous fluorescent agent within the patient's tissue as the patient's kidneys eliminate the exogenous fluorescent agent. Although the systems and apparatus disclosed below are described in the context of methods and apparatus for monitoring renal function, it should be understood that the disclosed systems and methods may be applied to any system and method for monitoring time-varying fluorescence emitted from a fluorescent agent within a diffusely reflective medium, where the optical properties of the diffusely reflective medium may also vary over time.

[0057] Figure 1 is a schematic diagram of a system 100 provided as a non-limiting example, wherein a device configured to detect only light having an emission wavelength (λ em ) of the photodetector 110, which detects photons having an emission wavelength (λ) from the region of interest of the patient 104. em Typically, the exogenous fluorescent agent 112 generates fluorescence 102 in response to an excitation event, including but not limited to: ex ) is illuminated by light 106, the occurrence of an enzyme reaction, a change in local potential, and any other known excitation event associated with an exogenous fluorescent agent. In one aspect, the system 100 may include a light source 108 configured to emit an excitation wavelength (λ ex ) is transmitted to the patient 104. In this regard, fluorescence 102 is generated in response to the illumination of the light 106. In addition, the excitation wavelength (λ ex ) and the emission wavelength of the fluorescent light 102 (λ em ) are spectrally different (i.e., λ ex and λ emcompletely different), such that the light detector 110 may be configured to selectively detect only the fluorescent light 102 by including any known light wavelength separation device, including but not limited to an optical filter.

[0058] In some aspects, changes in fluorescence 102 can be monitored to obtain information about the patient's physiological function or status. As a non-limiting example, a time-dependent decrease in fluorescence 102 measured after introducing an exogenous fluorescent agent 112 into the circulation of patient 104 can be analyzed to obtain information about the renal function of patient 104. In this non-limiting example, the rate of decrease in fluorescence 102 can be assumed to be proportional to the rate at which the patient's 104 kidneys eliminate the exogenous fluorescent agent 112, thereby providing measures of renal function, including but not limited to: renal decay time constant (RDTC) and glomerular filtration rate (GFR).

[0059] Without being limited to any particular theory, the intensity of the fluorescent light 102 detected by the light detector 110 may be affected by any one or more of a variety of factors, including but not limited to: ex The intensity or power of the light 106 transmitted to the patient 104 at λ, the scattering and absorption of the light 106 passing through the intervening tissue 114 of the patient 104 between the light source 108 and the exogenous fluorescent agent 112, the concentration of the exogenous fluorescent agent 112 illuminated by the light 106, the em The fluorescence 102 passing through the intervening tissue 114 of the patient 104 between the exogenous fluorescent agent 112 and the light detector 110 is scattered and absorbed, the excitation light 106 is transmitted by the light detector 110 configured to emit only at the wavelength λ em The leak-through of any optical filters that transmit light at this location, as well as the fluorescent light emitted by endogenous skin components.

[0060] Figure 14 is shown at the emission wavelength λ em Graph of the representative time history of the raw fluorescence signal obtained at the emission wavelength λ em The wavelength corresponds to the fluorescence emitted by the endogenous fluorescent agent in the patient's tissue in response to the illumination of the excitation wavelength light. The measured raw fluorescence signal (i.e., background signal 1402) obtained before the injection of the endogenous fluorescent agent may include autofluorescence (F auto ) and leakage of the excitation wavelength light (ExLT) through any optical filter configured to transmit only the emission wavelength light to the photodetector that produces the raw fluorescence signal. meas 1404) after the measured raw fluorescence signal may include the fluorescence signal superimposed on the background signal 1402 (ie, F auto and ExLT) of endogenous fluorescent agents (F agent )The intensity of the emitted fluorescence.

[0061] Existing methods generally assume that the optical properties within the intervening tissue 114 remain substantially constant throughout the period over which the system 100 obtains measurements. As a result, existing methods generally obtain initial measurements through the intervening tissue 114 of the patient 104 prior to the introduction of the exogenous fluorescent agent 112, and subtract these initial measurements to correct all subsequent data obtained after the introduction of the exogenous fluorescent agent 112. However, during long-term monitoring of the patient 104, changes in the optical properties of the intervening tissue 114 may occur due to changes in at least one characteristic, including but not limited to: the efficiency of light coupling from the photodetector 110 to the patient 104; the concentration of chromophores, such as hemoglobin, due to changes in blood volume caused by vascular dilation, constriction, or compression; changes in the optical properties of chromophores, such as hemoglobin, due to changes in oxygenation state; and changes in tissue structure, such as changes associated with edema.

[0062] Figure 16 is a graph of raw fluorescence signals measured before and after injection of an endogenous fluorescent agent, illustrating that the background signal can change over extended data acquisition periods associated with renal clearance of the endogenous fluorescent agent from a patient. Figure 16 As shown, the initial background signal level 1602 is approximately 0.01 intensity units higher than the final background signal level 1604, which was measured approximately 9 hours after the initial background signal level 1602. Without being limited to any particular theory, it is believed that the administration of blood pressure medication during the data acquisition period may have caused skin flushing and associated vasodilation of skin capillaries, which may have altered the optical properties of the patient's skin due to the increased concentration of blood, which contains hemoglobin, a known endogenous chromophore capable of absorbing light at both the excitation and emission wavelengths.

[0063] These dynamic changes in the optical properties of intervening tissue 114 can introduce uncertainty into long-term measurements of fluorescence 102. As a non-limiting example, changes in the optical properties of intervening tissue 114 can modulate the intensity or power of light 106 illuminating exogenous fluorescing agent 112, resulting in a modulation of fluorescence 102 generated by exogenous fluorescing agent 112, which can be erroneously interpreted as a modulation of the concentration of exogenous fluorescing agent 112. As another non-limiting example, changes in the optical properties of intervening tissue 114 can modulate the intensity or power of fluorescence 102 reaching light detector 110, which can also be erroneously interpreted as a modulation of the concentration of exogenous fluorescing agent 112. The potential modulation of changes in the optical properties of intervening tissue 114 can introduce uncertainty into measurements of fluorescence 102, particularly those associated with long-term monitoring of fluorescence 102, as described above.

[0064] Similarly, because of the autofluorescence (Fauto ) occurs in a similar manner to the fluorescence produced by exogenous fluorescent agents, so during the long-term measurement of fluorescence 102, dynamic changes in the optical properties of the intervening tissue may cause autofluorescence (F auto ) levels. As a non-limiting example, changes in the scattering and absorption of light 106 passing through intervening tissue 114 can modulate the intensity or power of light 106 illuminating endogenous chromophores, causing modulation of autofluorescence that can modulate background fluorescence during data acquisition. As another non-limiting example, changes in the scattering and absorption of autofluorescence passing through intervening tissue 114 can modulate the intensity of autofluorescence detected by light detector 110, which can modulate background fluorescence during data acquisition. Failure to properly account for potential modulation of background fluorescence can introduce uncertainty into raw fluorescence measurements and, by extension, into parameters derived from analysis of these fluorescence measurements.

[0065] As a non-limiting example, changes in autofluorescence associated with dynamic changes in the optical properties of a patient's skin can introduce uncertainty into the calculation of the renal decay time constant (RDTC), a measure of renal function as described below. Figure 15 Injection includes autofluorescence (IF Agent+AutoFlr , blue line) is a graph of the raw fluorescence signals measured before and after the addition of an endogenous fluorescer. Figure 15 The graph also includes the corrected fluorescence signal (IF Agent , green line), the corrected fluorescence signal is calculated by eliminating the effect of autofluorescence from the raw fluorescence signal using the method described below. Superimposed on each signal is the curve fit associated with the calculation of RDTC. Figure 15 As shown, the RDTC value of 2.76 h calculated using the raw fluorescence signal is much higher than the corresponding RDTC value of 2.31 h calculated using the corrected fluorescence signal.

[0066] In various aspects, a method is provided for correcting in vivo, real-time measurements of fluorescence from an exogenous fluorescent agent to eliminate the effects of changes in optical properties within a patient's tissue. Including an additional measurement of light that passes through the patient's tissue via an optical path separate from that of the fluorescence measurement (i.e., diffuse reflection) enhances the quantification of changes in the optical properties of the tissue during long-term monitoring of fluorescence from the exogenous fluorescent agent within the patient. Including this additional measurement in the correction method has been found in various aspects to significantly enhance the fidelity of the fluorescence measurement.

[0067] Provided herein below are detailed descriptions of apparatus for monitoring the fluorescence of exogenous fluorescent agents in vivo and methods for correcting fluorescence measurements to eliminate the effects of variations in background signal.

[0068] Although the devices and methods described below are described in the context of a non-invasive optical renal function monitor, it should be understood that the calibration methods described herein can be applied, with appropriate modifications, to any compatible device configured to perform measurements by transmitting EM radiation from an external source through any scattering medium and / or receiving EM radiation propagated through any scattering medium to an external detector. Non-limiting examples of EM radiation include visible light, near-IR light, IR light, UV radiation, and microwave radiation. The scattering medium can include any animate or inanimate material capable of transmitting EM radiation of at least one EM frequency without restriction. At least a portion of the scattering medium can also include one or more substructures or compounds capable of reflecting and / or absorbing EM radiation. Non-limiting examples of scattering media include: tissue of living or dead organisms, such as mammalian skin; gases, such as air, with or without attached particles (such as dust, liquid droplets, or solid particulate material); and fluids, such as water, with or without attached particles (such as air bubbles or solid particulate material). Furthermore, the devices and methods described below are not limited to the detection of renal function, but can be modified for use in detecting the function of other physiological systems, including, but not limited to, the liver or gastrointestinal systems.

[0069] System Description

[0070] In various aspects, the methods of correcting fluorescence measurements to eliminate the effects of changes in local skin properties as disclosed herein can be incorporated into any fluorescence monitoring system, including, but not limited to, systems for real-time and optical monitoring of renal function in vivo by measuring changes in fluorescence of an exogenous fluorescent agent injected into a patient as the agent is re-eliminated from the patient's body. Figure 2 1 is a block diagram of a system 200 for optically monitoring renal function of a patient 202 by fluorescence measurement of an exogenous fluorescent agent injected into the patient 202. The system 200 may include at least one sensor head 204 configured to transmit an excitation wavelength (λ ex ) is transmitted into a first region 206 of the patient 202. The system 200 is also configured to detect the emission wavelength (λ) at a second region 208 of the patient 202. em ) at the third region 210 of the patient 202, and detecting the excitation wavelength (λ ex ) and / or emission wavelength (λ em ) at the light.

[0071] System 200 may also include a controller 212, an operating unit 214, and a display unit 216 operably coupled to at least one sensor head 204. In various aspects, controller 212 is configured to control the operation of at least one sensor head 204, as described in further detail below. Controller 212 is also configured to receive measurements of light from at least one sensor head 204. Controller 212 is also configured to correct the light measurements for fluorescence from exogenous fluorescent agents according to at least one method, including but not limited to the disclosed method of correcting fluorescence measurements using measurements indicating dynamic changes in background signals associated with changes in autofluorescence and / or leakthrough of excitation wavelength light to a second light detector 224 configured to detect only emission wavelength light. Controller 212 is also configured to convert the fluorescence measurements received from at least one sensor head 204 into a summary parameter representing renal function of patient 202. Furthermore, controller 212 is configured to receive at least one signal representing user input from operating unit 214 and generate one or more displays on display unit 216, including but not limited to a graphical user interface (GUI).

[0072] A detailed description of the sensor head 204 and the controller 212 is provided below.

[0073] A.Sensor head

[0074] In various aspects, the sensor head 204 includes at least one light source and at least one light detector in a housing. Figure 6 6 is a side view of a housing 600 for a sensor head 204, comprising an upper housing 602 and a lower housing 604 connected together to enclose two light sources and two light detectors. The bottom surface 608 of the lower housing 604 also includes a contact surface 606 that is configured to be attached to the skin of the patient 202 using a biocompatible adhesive material (including but not limited to a surgical adhesive). In use, the surface of the adhesive material opposite the contact surface 606 can be secured to the skin of the patient 202. In various aspects, the adhesive material can be configured to transmit light through the light source into the patient and also transmit fluorescence from the patient to the light detector. In one aspect, the adhesive material can be an optically transparent material. In another aspect, the adhesive material can be made of a non-fluorescent material to prevent the adhesive material from generating contaminating fluorescence.

[0075] In various other aspects, the upper housing 602 may also include one or more openings 806 configured to provide access to the interior of a cable (including but not limited to a USB cable) and / or to provide a window for a display (such as an indicator LED) generated by circuitry contained within the housing 600.

[0076] Figure 7 yes Figure 86. A bottom view of the housing 600 is shown. The contact surface 606 can include an aperture plate 702 including one or more apertures 704 configured to transmit light between the patient's skin and a light source and light detector housed within the housing 600. In one aspect, the aperture plate 702 can be epoxied into the lower housing 604 to prevent liquid from entering the interior of the housing 600. In various aspects, the size, arrangement, and / or spacing of the one or more apertures 704 can be selected to enhance various aspects of the operation of the system 200, as described in further detail below. In another aspect, the contact surface 606 can also include a temperature sensor opening 706 configured to provide a thermal path from the patient's skin surface to an additional temperature sensor 228 configured to monitor the temperature at the patient's skin surface.

[0077] Figure 8 Schematic diagram showing the arrangement of electrical components within the housing 600. Figure 8 The upper shell 602 and the lower shell 604 can be fixed together with screws 802, and the screw holes and the interface between the two shell members can be filled with a waterproof filling material 804, which includes but is not limited to silicone materials, such as room temperature vulcanized silicone (RTV), to prevent liquid from entering the interior of the shell 600.

[0078] In one aspect, the housing 600 can further include a cable opening 806 formed through the upper housing 602. The cable opening 806 can be configured to provide access to the interior of a cable, including but not limited to a USB cable. In one aspect, the cable can be capable of providing power to the light source, light detectors, indicator lights, and associated electrical devices and circuits, as described below. In another aspect, the cable can also enable control signals to be transmitted into the housing to enable operation of electrical components within the housing 600, and the cable can also enable transmission of data signals that encode measurements obtained by one or more of the sensor devices contained within the housing 600, the sensor devices including but not limited to: the first light detector 222, the second light detector 224, any additional light detectors (such as the first monitor photodiode 904 and the second monitor diode 906), and any additional temperature sensor 228 (see Figure 9 In one aspect, the cable can be attached to the cable opening 806 and the adjacent upper housing 602 with a light absorbing adhesive, including but not limited to black epoxy, and can also be sealed with a waterproof filler material, including but not limited to RTV, to prevent water intrusion.

[0079] In another aspect, the housing 600 can further include at least one display opening 808 formed through the upper housing 602. In one aspect, each display opening 808 can be configured to provide a window for a display (such as an indicator LED 810) generated by circuitry contained within the housing 600. In one aspect, each indicator LED 810 can be positioned on a circuit board 812. In one aspect, a light pipe 814 can be epoxied into the display opening 808 within the upper housing 602 above each indicator LED 810. Each light pipe 814 can be filled with a waterproof filler material, such as RTV, for liquid ingress protection. In various aspects, the at least one indicator LED 810 can illuminate in a predetermined pattern to enable a user of the system 200 to monitor the operating status of the sensor head 204.

[0080] Figure 9 is a close-up view of the internal optical region of the sensor head 204, illustrating, in one aspect, the arrangement of the light sources 218 / 220 and light detectors 222 / 224 within the housing 600. In one aspect, the light sources 218 / 220 are separated from the light detectors 222 / 224, and the first light detector 222 and the second light detector 224 are separated from each other by a sensor mount 912 secured to the aperture plate 702. In one aspect, the sensor mount 912 ensures that light from the light sources 218 / 220 does not reach the light detectors 222 / 224 without being coupled through the skin of the patient 202. The separation between the first light detector 222 within the first detection well 908 and the second light detector 224 within the second detection well 910 ensures that the fluorescence signal generated by exogenous fluorescing agents within the tissue of the patient 202 is distinguishable from the unfiltered excitation light introduced by the first light source 218.

[0081] Refer again Figure 9 , the sensor mount 912 can be aligned with a circuit board (not shown) containing the light sources 218 / 220 and the light detectors 222 / 224 using alignment pins 914, and the sensor mount 912 can be held in place using screws 916. In one aspect, the sensor mount 912 can be secured to the circuit board containing the light sources 218 / 220 and the light detectors 222 / 224 using a light absorbing adhesive (including but not limited to black epoxy). In this regard, this light resistant bond between the circuit board and the sensor mount 912 inhibits light leakage between the light sources 218 / 220 and the light detectors 222 / 224, and also inhibits light leakage between the first light detector 222 and the second light detector 224. The aperture 704 configured to transmit light to and from the skin beneath the contact surface 606 of the sensor head 204 is formed by a structurally separate aperture plate 702 (see Figure 7) are formed to provide precise alignment of the aperture 704 with the corresponding light source 218 / 220 and light detector 222 / 224, as will be described in more detail below.

[0082] In various aspects, the sensor mount 912 can also provide electrical shielding for any sensitive electrical devices within the sensor head 204, including but not limited to the light detectors 222 / 224. In one aspect, the sensor mount 912 can be constructed of a conductive material, including but not limited to aluminum and aluminum alloys. In this regard, the sensor mount 912 can be electrically coupled to the circuit board ground using a conductive screw 916. Additionally, any glass windows within the source well 902 and / or detector wells 908 / 910 adjacent to the aperture plate 702, including but not limited to the optical filter 244 and the clear glass 246 (see FIG. 2 ), can be electrically coupled to the circuit board ground using a conductive screw 916. Figure 2 )), may further include a conductive coating. Non-limiting examples of suitable conductive coatings for the glass window of the sensor mount include conductive indium tin oxide (ITO) coatings and any other suitable transparent and conductive coatings.

[0083] Without being limited to any particular theory, the conductive material of sensor mount 912 provides a partial Faraday cage to shield electrically sensitive detectors 222 / 224 from electrical noise generated or conducted by the patient's body. The partial Faraday cage provided by sensor mount 912 can be completed with a conductive ITO coating on the glass windows within source well 902 and / or detector wells 908 / 910. In one aspect, the conductive coating on the glass windows, such as the ITO coating, is sufficiently conductive to provide electrical shielding while remaining sufficiently transparent to transmit light to and from the skin surface of patient 202. In another aspect, the ITO coating on each glass window can be grounded to the conductive sensor mount 912 using any known electrical grounding method, including but not limited to: connecting the glass coating to a wire connected to sensor mount 912, with sensor mount 912 attached at both ends with conductive epoxy, or directly attaching the coated glass to a glass fitting, such as a flange or frame formed within each source well 902 and / or detector well 908 / 910, using conductive epoxy.

[0084] In various aspects, the contact surface 606 of the housing 600 can be attached to the patient's skin using a biocompatible and adhesive material 610 (including but not limited to a clear double-sided medical grade adhesive), such as Figure 6 and Figure 7As described herein, any adhesive material is selected to be optically transmissive at the excitation and emission wavelengths used by the system 100. The adhesive material 610 can be positioned on the contact surface 606 such that the adhesive material covers the aperture 704 but exposes the temperature sensor opening 706 to ensure sufficient thermal contact with the skin of the patient 202. In one aspect, the sensor head 204 can be further secured to the patient 202 using one or more additional biocompatible medical fastener devices, including but not limited to Tegaderm bandages, medical tape, or any other suitable biocompatible medical fastener devices, as desired.

[0085] In one aspect, the contact surface 606 can be located near the leading edge of the sensor head 204 to provide for accurate positioning of the contact surface 606 on a selected area of the patient's skin. In another aspect, the apertures 704 can be located toward the center of the contact surface 606 to reduce the ingress of ambient light. Without being limited to any particular theory, ambient light can enter one or more of the apertures 704 due to incomplete adhesion of the contact surface 606 to the patient's skin and / or due to ambient light propagating through exposed patient skin just outside the footprint of the contact surface 606 into the apertures 704.

[0086] Refer again Figure 6 , the bottom surface 608 of the sensor head 204 is curved away from the plane of the contact surface 606 to enable the sensor head 204 to be attached to different body types and locations. To attach the sensor head 204 to a relatively flat or concave surface, any gap 612 between the bottom surface 608 and the skin surface of the patient 202 can be filled with biocompatible foam to ensure consistent contact with the patient 202.

[0087] i) Light source

[0088] In various aspects, each sensor head 204 includes a first light source 218 and a second light source 220 configured to transmit light to a first region 206 of a patient 202. The first light source 218 is configured to transmit light at an excitation wavelength, and the second light source 220 is configured to transmit light at an emission wavelength. In one aspect, the excitation wavelength can be selected to fall within a spectral range where the exogenous fluorescing agent exhibits relatively high absorption. In another aspect, the emission wavelength can be selected to fall within a spectral range where the exogenous fluorescing agent exhibits relatively high emission. The exogenous fluorescing agent can be selected to enhance contrast relative to other chromophores within the tissue of the patient 202, including, but not limited to, hemoglobin within red blood cells and / or melanin within melanocytes. In various aspects, the exogenous fluorescing agent can be selected to measure within a spectral range where the absorption of other chromophores, such as hemoglobin, within the tissue of the patient 202 varies less during use.

[0089] Without being limited to any particular theory, hemoglobin (Hb) is an absorber of visible light in the tissues of patient 202, and if the Hb absorber changes during the measurement period of system 200, it has the potential to interfere with the fluorescence measurement of the exogenous fluorophore. Because hemoglobin (Hb) is capable of gas exchange in almost all tissues containing circulatory blood vessels, almost all tissues are susceptible to interference with the fluorescence measurement of system 200 due to fluctuations in hemoglobin concentration. In most tissues, externally applied pressure can cause blood pooling, which can manifest as an apparent attenuation of fluorescence measured at the skin surface. The periodic opening and closing of blood vessels near the skin surface ("vasomotion") can also cause fluctuations in hemoglobin concentration, which can introduce additional noise into the fluorescence measurement of the exogenous fluorophore by system 200. In addition, in some patients 202, such as those with lung disease, changes in the oxygenation state of Hb can also be observed, due to differences in the absorption spectra of deoxygenated hemoglobin (Hb) and oxyhemoglobin (HbO2), resulting in additional potential changes in background skin absorption, such as Figure 3 shown.

[0090] In one aspect, the excitation and emission wavelengths for the exogenous fluorescer can be selected to coincide with a pair of HbO2 / Hb isosbestic points, each of which, as defined herein, is a wavelength characterized by approximately equal light absorbance by HbO2 and Hb. Without being limited to any particular theory, as long as the combined concentration of HbO2 and Hb remains relatively stable during fluorescence measurements by system 200, fluorescence measurements taken at each isosbestic wavelength are less sensitive to changes due to changes in hemoglobin oxygenation. Non-limiting examples of Hb / HbO2 isosbestic wavelengths include approximately 390 nm, approximately 422 nm, approximately 452 nm, approximately 500 nm, approximately 530 nm, approximately 538 nm, approximately 545 nm, approximately 570 nm, approximately 584 nm, approximately 617 nm, approximately 621 nm, approximately 653 nm, and approximately 805 nm.

[0091] In various aspects, the excitation wavelength and emission wavelength can be selected based on the absorption wavelength and emission wavelength of the selected exogenous fluorescer of system 200. In one aspect, the excitation wavelength can be an isoabsorbent wavelength of HbO2 / Hb and simultaneously a wavelength within the high absorbance spectral range of the exogenous fluorescer. In another aspect, the emission wavelength can be an isoabsorbent wavelength of HbO2 / Hb and simultaneously a wavelength within the spectral range emitted by the exogenous fluorescer. Table 1 provides a summary of isoabsorbent wavelengths of HbO2 / Hb within the spectral range of 200 nm to approximately 1000 nm. Figure 4 This is a graph showing the absorption spectra of HbO2 / Hb and other absorption wavelengths shown in Table 1.

[0092] Table 1. HbO2 / Hb isosbesty wavelength λ = 200 to 1000 nm

[0093]

[0094] As an illustrative example, Figure 3 The graph summarizes the absorption spectra of HbO2 and Hb, and the frequency spectrum of the exogenous fluorescent agent MB-102. The emission spectra of the blue LED light source and the green LED light source are also shown superimposed on Figure 3 In this aspect, the system 200 may include a blue LED as the first light source 218, and the excitation wavelength of the system 200 may be an isosbestic wavelength of about 450 nm. As listed in Table 1 and Figure 3 As shown, the Hb absorbance spectrum is significantly tilted at the isosbestic wavelength of approximately 420 nm to approximately 450 nm (see columns 3 and 4 of Table 1), indicating that the relative absorbance of HbO2 and Hb at the isosbestic wavelength of approximately 450 nm is sensitive to small changes in the excitation wavelength. However, at wavelengths above approximately 500 nm, the HbO2 / Hb spectrum is less steeply tilted, and a wider bandwidth light source (including but not limited to an LED with a bandpass filter) may be sufficient for use as the first light source 218.

[0095] In another aspect, the excitation wavelength can be selected to enhance the light absorption contrast between the exogenous fluorescent agent and the chromophore within the tissue of the patient 202. As a non-limiting example, Figure 3 As shown, at an isosbestic wavelength of 452 nm, MB-102 absorbs more than three times more light than HbO2 and Hb. Without being limited to any particular theory, MB-102 absorbs a higher proportion of light irradiating the tissue of patient 202 at a wavelength of approximately 450 nm relative to HbO2 and Hb, thereby enhancing MB-102's absorption efficiency and reducing the intensity of light at the excitation wavelength required to induce a detectable fluorescent signal.

[0096] In various aspects, a second isosbestic wavelength may also be selected as the emission wavelength of the system 200. As a non-limiting example, Figure 3The emission spectrum of the MB-102 exogenous contrast agent is shown, which is characterized by an emission peak at a wavelength of approximately 550 nm. In this non-limiting example, an isosbestic wavelength of 570 nm can be selected as the emission wavelength to be detected by the first detector 222 / 224 and the second detector 222 / 224. In various other aspects, the emission wavelength of the system 200 can be selected to fall within a spectral range characterized by a relatively low absorbance of chromophores within the tissue of the patient 202. Without being limited to any particular theory, the low absorbance of the chromophore at the selected emission wavelength can reduce the loss of light emitted by the exogenous fluorescent agent and improve the efficiency of fluorescence detection.

[0097] In various aspects, the first light source 218 and the second light source 220 can be any light source configured to transmit light at an excitation wavelength and an emission wavelength. Typically, the first light source 218 transmits light to the exogenous fluorescing agent at an intensity sufficient to penetrate the tissue of the patient 202, wherein sufficient intensity remains to induce the exogenous fluorescing agent to emit light at the emission wavelength. Typically, the first light source 218 transmits light to the exogenous fluorescing agent at an intensity sufficient to penetrate the tissue of the patient 202, wherein the intensity remaining after scattering and / or absorption is sufficient to induce fluorescence at the emission wavelength by the exogenous fluorescing agent. However, the intensity of the light transmitted by the first light source 218 is limited to an upper limit to prevent adverse effects of exogenous fluorescing agents and / or endogenous chromophores ("autofluorescence") in the skin, such as tissue burning, cell damage, and / or photobleaching.

[0098] Similarly, second light source 220 transmits light at the emission wavelength of the exogenous fluorescer at an intensity configured to provide sufficient energy to propagate through first region 206 of the patient by scattering and absorption and to propagate with sufficient residual intensity out of second region 208 and third region 210 for detection by first light detector 222 and second light detector 224, respectively. As with first light source 218, the intensity of the light generated by second light source 220 is limited to an upper limit to prevent adverse effects such as tissue damage or photobleaching as previously described.

[0099] In various aspects, the first light source 218 and the second light source 220 can be any light source suitable for use with fluorescence medical imaging systems and devices. Non-limiting examples of suitable light sources include: LEDs, diode lasers, pulsed lasers, continuously oscillating lasers, xenon arc lamps or mercury vapor lamps with excitation filters, lasers, and supercontinuum light sources. In one aspect, the first light source 218 and / or the second light source 220 can generate light with a narrow spectral bandwidth suitable for monitoring the concentration of exogenous fluorescent agents using the methods described herein. In another aspect, the first light source 218 and the second light source 220 can generate light with a relatively wide spectral bandwidth.

[0100] In one aspect, the selection of the intensity of light generated by the first light source 218 and the second light source 220 of the system 200 can be influenced by any one or more of at least several factors, including, but not limited to, the maximum permissible exposure (MPE) for skin exposure to the laser beam according to applicable regulatory standards such as ANSI standard Z136.1. In another aspect, the light intensity used in the system 200 can be selected to reduce the potential for photobleaching of exogenous fluorescent sources and / or other chromophores within the tissue of the patient 202, including, but not limited to, collagen, keratin, elastin, hemoglobin within red blood cells, and / or melanin within melanocytes. In yet another aspect, the light intensity used in the system 200 can be selected to elicit a detectable fluorescence signal from exogenous fluorescent sources within the tissue of the patient 202 and within the first light detector 222 and / or the second light detector. In yet another aspect, the light intensity used in the system 200 can be selected to provide appropriately high light energy while reducing power consumption, inhibiting heating / overheating of the first light source 218 and the second light source 220, and / or reducing the time the patient's skin is exposed to light from the first light detector 222 and / or the second light detector.

[0101] In various aspects, the intensity of the first and second light sources 218, 220 can be modulated to compensate for any one or more of at least several factors, including, but not limited to, individual differences in chromophore concentration within patient 202, such as variations in skin pigmentation. In various other aspects, the detection gain of the light detector can be modulated to similarly compensate for variations in individual differences in skin properties. In one aspect, variations in skin pigmentation can occur between two different individual patients 202, or between two different locations on the same patient 202. In one aspect, light modulation can compensate for variations in the optical path taken by light through the tissue of patient 202. The optical path can vary due to any one or more of at least several factors, including, but not limited to, variations in the separation distance between the light source and light detector of system 200; variations in the secure attachment of sensor head 204 to the skin of patient 202; variations in the light output of the light source due to exposure of the light source to environmental factors such as heat and humidity; variations in the sensitivity of the light detector due to exposure of the light detector to environmental factors such as heat and humidity; modulation of the duration of illumination by the light source; and any other relevant operational parameters.

[0102] In various aspects, the first light source 218 and the second light source 220 can be configured to modulate the intensity of the light produced as needed based on any one or more of the factors described above. In one aspect, if the first light source 218 and the second light source 220 are devices configured to continuously change the output flux as needed, such as LED light sources, the intensity of the light can be electronically modulated using methods including, but not limited to, modulating the potential, current, and / or power supplied to the first light source 218 and / or the second light source 220. In another aspect, the intensity of the light can be modulated using optical methods including, but not limited to, using optical devices including, but not limited to, an iris, a shutter, and / or one or more filters to partially or completely block light exiting the first light source 218 and the second light source 220; using optical devices including, but not limited to, lenses, reflectors, and / or prisms to divert the path of light exiting the first light source 218 and the second light source 220 away from the first area 206 of the patient.

[0103] In various aspects, the intensity of the light generated by the first light source 218 and the second light source 220 can be modulated via control of the laser flux, which is defined herein as the energy rate within the generated light beam. In one aspect, the laser flux can be limited to a range defined by safety standards, including but not limited to ANSI standards for exposure to laser energy such as ANSI Z136.1. Without being limited to any particular theory, the maximum light flux delivered to the patient 202 can be affected by a variety of factors, including but not limited to the wavelength of the light delivered and the duration of exposure to the light. In various aspects, the maximum light flux ranges from about 0.003 J / cm2 for light delivered at a wavelength less than about 302 nm to about 1 J / cm2 for light delivered at a wavelength in the range of about 1500 nm to about 1800 nm, for a duration of up to about 10 seconds. For light delivered in the wavelength range of about 400 nm to about 1400 nm (visible / NIR light), the maximum flux can be about 0.6 J / cm² for a duration of up to about 10 seconds, and about 0.2 J / cm² for a duration in the range of about 10 seconds to about 30,000 seconds. For extended exposure, the delivered light is limited to a maximum power density (W / cm²) according to ANSI standards: visible / NIR light is limited to 0.2 W / cm², and far IR light is limited to about 0.1 W / cm². Without being limited to a particular theory, extended exposure to light delivered at UV wavelengths is generally not recommended according to ANSI standards.

[0104] In another aspect, the light flux at the excitation wavelength generated by the first light source 218 can be modulated to provide sufficient energy to propagate through the skin in the first region 206 of the patient 202 to the exogenous fluorescent agent without photobleaching, and to illuminate the exogenous fluorescent agent with sufficient energy to induce detectable fluorescence at the first light detector 222 and / or the second light detector 224. In another aspect, the light flux at the emission wavelength generated by the second light source 220 can be modulated to provide sufficient energy to propagate through the skin in the first region 206 of the patient 202 and through the skin in the second region 208 and the third region 210 without photobleaching to appear as detectable light at the first light detector 222 and the second light detector 224, respectively. As a non-limiting example, the light flux generated by the light source at 450 nm or 500 nm can be limited to 1.5 mW / cm, respectively. 2 and 5mW / cm 2 to prevent photobleaching.

[0105] In various aspects, without limitation as described above, the light flux generated by the first light source 218 and the second light source 220 can be modulated by any suitable system and / or device. The modulation can be enabled once during operation of the system 200, and thus, the light flux generated by each of the first light source 218 and the second light source 220 can be relatively constant throughout operation of the system 200. In another aspect, the light modulation can be enabled at discrete times during operation of the system 200, or the light modulation can be enabled continuously during operation of the system 200.

[0106] In one aspect, when the system 200 is configured in the engineering mode, the light flux can be modulated via manual adjustment of any power settings and / or optical device settings as described above. In another aspect, the light flux can be automatically modulated via one or more control schemes encoded in the light source control unit of the controller 212, as described below. In this aspect, the degree of modulation can be specified at least in part based on feedback measurements obtained by various sensors disposed in the sensor head 204 of the system 200, including but not limited to the additional light detector 226 and the temperature sensor 228, as described in more detail below.

[0107] In various aspects, the light generated by the first light source 218 and the second light source 220 is further characterized by a pulse width, defined herein as the duration of the generated light. While pulse width is often used to characterize the performance of a light source that generates light in discrete pulses (such as a pulsed laser), it should be understood that the term "light pulse" as used herein refers to any discrete light pulse generated by a single light source at a single wavelength to enable acquisition of a single fluorescence measurement by the system 200. Similarly, the term "pulse width" as used herein refers to the duration of a single light pulse generated by a single light source. The pulse width is typically selected based on one or more of at least several factors, including, but not limited to: delivering sufficient light energy to induce detectable fluorescence from exogenous fluorescers or other chromophores within the tissue of the patient 202 without photobleaching; adhering to safety standards, such as ANSI standards, for the light delivered to the patient; delivering light at a sufficiently high rate to enable data acquisition at a rate compatible with real-time monitoring of renal function; the performance capabilities of the selected light source, light detector, and other devices of the system 200; preserving the useful life of the light source, light detector, and other devices associated with generating and detecting light energy; and other relevant factors.

[0108] In various aspects, the pulse width of light generated by the first light source 218 and the second light source 220 can be independently selected to have a duration ranging from about 0.0001 seconds to about 0.5 seconds. In various other aspects, the pulse width of light generated by the first light source 218 and the second light source 220 can be independently selected to have a duration ranging from about 0.0001 seconds to about 0.001 seconds, from about 0.0005 seconds to about 0.005 seconds, from about 0.001 seconds to about 0.010 seconds, from about 0.005 seconds to about 0.05 seconds, from about 0.01 seconds to about 0.1 seconds, from about 0.05 seconds to about 0.15 seconds, from about 0.1 seconds to about 0.2 seconds, from about 0.15 seconds to about 0.25 seconds, from about 0.2 seconds to about 0.3 seconds, from about 0.25 seconds to about 0.35 seconds, from about 0.3 seconds to about 0.4 seconds, from about 0.35 seconds to about 0.45 seconds, and from about 0.4 seconds to about 0.5 seconds. In one aspect, the pulse widths of the light generated by the first light source 218 and the second light source 220 are both about 0.1 seconds, e.g. Figure 5 Schematically shown in .

[0109] In another aspect, the light generated by the first light source 218 and the second light source 220 can also be characterized by a pulse rate, which is defined herein as the number of pulses generated by the light source per second. Although pulse rate is often used to characterize the performance of a light source that generates discrete pulses of light, such as a pulsed laser, it should be understood that the term "pulse rate" as used herein refers to the rate at which a single light source generates discrete light pulses at a single wavelength, which is associated with the acquisition of fluorescence measurements by the system 200. In various aspects, the pulse rate can be selected based on one or more of at least several factors, including but not limited to: compliance with safety standards, such as ANSI standards, for delivering light to the patient; performance capabilities of the selected light sources, light detectors, and other devices of the system 200; compatibility of the rate at which light is delivered with a data acquisition rate that is fast enough for real-time monitoring of renal function; preservation of the useful life of the light sources, light detectors, and other devices associated with generating and detecting light energy; and any other relevant factors.

[0110] In various aspects, the light source is configured to deliver light into the tissue of the patient 202 at a single location, such as the first region 206. Figure 2 In one aspect, delivering light at both the excitation wavelength and the emission wavelength to the same first region 206 enables the two light pulses to share at least a portion of the optical path through the tissue of the patient 202 between an entry point at the first region 206 and detection points at the second region 208 and the third region 210. As discussed in detail below, this arrangement of the optical paths enhances the quality of the data generated by the system 200.

[0111] In one aspect, the first light source 218 and the second light source 220 are operably coupled to a common means of light delivery. In one aspect (not shown), each of the first light source 218 and the second light source 220 can be operably coupled to a first optical fiber and a second optical fiber, respectively, and the first optical fiber and the second optical fiber can be connected to a third optical fiber that is configured to direct light from the first optical fiber and / or the second optical fiber into the first region 206 of the patient 202. In another aspect, the first light source 218 and the second light source 220 can be operably coupled to a common optical fiber or other optical component that is configured to direct light from the first light source 218 and / or the second light source 220 into the first region 206 of the patient 202. In this aspect, the light generated by the first light source 218 and the second light source 220 can be directed into the common optical fiber or other optical component in an alternating pattern using an adjustable optical device, including but not limited to a dichroic mirror or a rotating mirror.

[0112] In one aspect, the system 200 may include a sensor head 204 having a sensor mount 912 configured with one or more wells into which light sources 218 / 220 and light detectors 222 / 224 may be attached in a predetermined arrangement. Figure 9 and Figure 10 As shown, the first light source 218 and the second light source 220 can be located within a source well 902 of a sensor mount 912 positioned within the sensor head 204 (see Figure 9 In one aspect, the source well 902 can include a first LED light source 218 generating light at an excitation wavelength and a second LED light source 220 generating light at an emission wavelength, the first LED light source 218 and the second LED light source 220 being operably coupled to a single light delivery aperture 1002 formed through the aperture plate 702 (see Figure 10 ), the single light delivery aperture ensures that the two wavelengths of light (i.e., excitation and emission) enter the skin of the patient 202 at approximately the same location (including but not limited to the first region 206), as shown in FIG. Figure 2 In one aspect, the source well 902 further comprises a first monitor photodiode 904 and a second monitor photodiode 906, which are used to correct for variations in the output power from the LED light source, as described in further detail below.

[0113] In one aspect, only a portion of the light energy generated by the LED light sources is transmitted to the skin of the patient 202 via a single light delivery aperture 1002. In one aspect, the skin of the patient 202 receives approximately 1% of the light energy generated by the LED light sources. In various other aspects, the skin of the patient 202 receives approximately 2%, approximately 3%, approximately 4%, approximately 5%, approximately 7.5%, approximately 10%, approximately 20%, and approximately 50% of the light energy generated by the LED light sources. Without being limited to any particular theory, the portion of light generated by the LED light sources that is transmitted to the skin of the patient 202 can be increased by incorporating additional optical elements configured to focus and / or direct light from each LED light source to the light delivery aperture 1002. In another aspect, a diffuser can be used to mix the output of the light sources so that the light energy is uniformly presented at the surface of the patient's skin.

[0114] ii) Photodetector

[0115] Refer again Figure 2, in various aspects the system 200 further comprises a first light detector 222 and a second light detector 224. In one aspect, the first light detector 222 is configured to measure unfiltered light emitted from the tissue of the patient 202 at the second region 208, and the second light detector 224 is configured to measure filtered light emitted from the tissue of the patient 202 at the third region 210. In this regard, the second light detector 224 further comprises an optical filter 244 that is configured to block light at the excitation wavelength. As a result, the first light detector 222 is configured to measure light received at both the excitation wavelength and the emission wavelength, and the second light detector 224 is configured to detect light received only at the emission wavelength. In combination with irradiating the tissue of the patient 202 with light only at the excitation wavelength and light only at the emission wavelength in an alternating series (see Figure 5 ), the measurements from the first light detector 222 and the second light detector 224 can be analyzed as described below to measure the fluorescence of the exogenous fluorescent agent and to correct the fluorescence measurements by eliminating the effects of dynamic changes in the background signal according to the correction method described below in this article.

[0116] In various aspects, light from a second region 208 and a third region 210 within the tissue of the patient 202 is detected by a first light detector 222 and a second light detector 224, respectively. The second region 208 and the third region 210 are each separated by a nominal distance from the first region 206. Light generated by the first light source 218 and the second light source 220 is transmitted to the first region 206. This nominal separation distance can be selected to balance two or more effects that may affect the quality of the data detected by the light detectors. Without being limited to any particular theory, as the nominal separation distance increases, the total detected signal from the light detectors may decrease due to light scattering along the longer optical path between the light source and the light detector. This effect can be mitigated by the choice of emission wavelength, which can result in a less significant reduction in the detected fluorescence signal (i.e., light at the emission wavelength) relative to the signal associated with light detected at the excitation wavelength as the nominal separation distance increases. A longer nominal separation distance results in greater sensitivity to signal changes due to changes in the optical properties of the tissue.

[0117] In one aspect, the nominal separation distance can be in the range of from 0 mm (i.e., co-location of the light source and light detector) to about 10 mm. In various other aspects, the nominal separation distance can be in the range of from about 1 mm to about 8 mm, from about 2 mm to about 6 mm, and from about 3 mm to about 5 mm. In various additional aspects, the nominal separation distance can be 0 mm, about 1 mm, about 2 mm, about 3 mm, about 4 mm, about 5 mm, about 6 mm, about 8 mm, and about 10 mm. In one aspect, the nominal separation distance can be about 4 mm to balance the competing effects of the logarithmic decrease in signal and the reduced size of the background signal relative to the signal from the exogenous fluorescent agent.

[0118] Refer again Figure 9 , a first light detector 222 can be positioned within the first detection well 908 of the sensor mount 912, and a second light detector 224 can be positioned within the second detection well 910 of the sensor mount 912 within the sensor head 204. The first light detector 222 and the second light detector 224 can receive light from the tissue of the patient 202 through the first detector aperture 1004 and the second detector aperture 1006, respectively. In one aspect, the first detector aperture 1004, the second detector aperture 1006, and the light transmission aperture 1002 are separated from each other by a nominal separation distance as disclosed above herein, including, but not limited to, a nominal separation distance of 4 mm. In one aspect, the first detection well 908, the second detection well 910, and the light source well 902 of the sensor mount 912 can be optically isolated from each other to ensure that light from the light source 218 / 220 does not reach the light detectors 222 / 224 without being coupled through the skin of the patient 202. As described in detail below, the separation between the two detection wells 908 / 910 ensures that the detected fluorescent signal from the exogenous fluorescer can be distinguished from the unfiltered excitation light.

[0119] In one aspect, the three holes 704 of the orifice plate 702 (see Figure 7 ) are circular and have a diameter ranging from about 0.5 mm to about 5 mm. In various other aspects, the diameter of the holes can range from about 0.5 mm to about 1.5 mm, about 1 mm to about 2 mm, about 1.5 mm to about 2.5 mm, about 2 mm to about 3 mm, about 2.5 mm to about 3.5 mm, about 3 mm to about 4 mm, about 3.5 mm to about 4.5 mm, and about 4 mm to about 5 mm.

[0120] In one aspect, the three apertures 704 of the aperture plate 702 are circular apertures having a diameter of approximately 1 mm. Because the logarithm of the signal decreases with increasing separation distance from the light source at the skin interface of the sensor head 204, this finite width of the apertures can result in an effective source-detector separation that is less than the nominal separation distance.

[0121] In various aspects, the light detectors 222 / 224 of the system 200 can be any suitable light detecting device, but are not limited thereto. Non-limiting examples of suitable light detecting devices include: photoemission detectors, such as photomultiplier tubes, phototubes, and microchannel plate detectors; photodetectors, such as reverse-biased LEDs used as photodiodes, photoresistors, photodiodes, phototransistors; and any other suitable light detecting device. In one aspect, the light detectors 222 / 224 are sufficiently sensitive to detect fluorescence emitted by exogenous fluorescent agents within the tissue of the patient 202, the tissue of the patient 202 including melanin in the epidermis ranging from about 1% to about 40% and blood volume ranging from about 0.5% to about 2% of the skin volume. In one aspect, the light detectors 222 / 224 can be silicon photomultiplier (SPM) devices.

[0122] In one aspect, the first light detector 222 can be configured to detect light at both the excitation frequency and the emission frequency, and the second light detector 224 can be configured to detect light only at the emission frequency. In one aspect, due to the design and materials of the sensor element of the second light detector 224, the second light detector 224 can respond only to light at the emission wavelength. In another aspect, the second light detector 224 can respond to a wider range of light wavelengths, but can be located downstream of an optical filter that is configured to pass only a portion of incident light having the emission wavelength and is also configured to block the passage of light having wavelengths other than the emission wavelength.

[0123] Any suitable optical filter can be selected for use with the second light detector 224 to selectively detect light at the emission wavelength. Non-limiting examples of suitable optical filters include absorption filters and interference / dichroic filters. Without being limited to any particular theory, the performance of absorption filters does not vary significantly with the angle of the incident light, whereas the performance of interference / dichroic filters is sensitive to the angle of the incident light and may require additional collimating optics to effectively filter the Lambertian light distribution representative of the light emitted from the skin of the patient 202.

[0124] In one aspect, the second light detector 224 can be located downstream of an absorptive longpass filter configured to pass light greater than a predetermined wavelength to the second light detector 224. As a non-limiting example, the second light detector 224 can be located downstream of a longpass OG 530 filter configured to pass light greater than approximately 530 nm. Other non-limiting examples of suitable filters include a Hoya O54 filter and a Hoya CM500 filter.

[0125] In various aspects, an optical filter 244 configured to absorb light at the excitation wavelength can be positioned within the second detection well 910 between the second light detector 224 and the second detector aperture 1006. In one aspect, the optical filter 244 can be constructed from OG 530 Schott glass. The thickness of the optical filter 244 can be selected to provide an optical density sufficient to filter the excitation light by approximately three orders of magnitude. In one aspect, the thickness of the optical filter 244 can be in the range of about 1 mm to about 10 mm. In various other aspects, the thickness of the optical filter 244 can be in the range of about 1 mm to about 8 mm, about 2 mm to about 6 mm, and about 3 mm to about 5 mm. In various additional aspects, the thickness of the optical filter 244 can be about 1 mm, about 2 mm, about 3 mm, about 4 mm, about 5 mm, about 6 mm, about 7 mm, about 8 mm, about 9 mm, and about 10 mm. In one aspect, the optical filter 244 is a 3-mm thick filter constructed from OG 530 Schott glass.

[0126] In another aspect, an optical diffuser can be disposed within the light source well 902. In this aspect, the optical diffuser can mix light from the first light source 218 / 220 and the second light source 218 / 220 that enter the light source well 902. By using the optical diffuser to mix the light from the first light source 218 / 220 and the second light source 218 / 220 prior to illuminating the first region 206 of the patient 202, the similarity of the optical paths taken by the emission wavelength light and the excitation wavelength light of the patient's tissue is enhanced relative to the corresponding optical paths taken by unmixed light, thereby reducing potential sources of variation.

[0127] In one aspect, a transparent material configured to pass light at both the excitation wavelength and the emission wavelength can be located within the first detection well 908 between the first light detector 222 and the first detector aperture 1004. In this regard, the transparent material can be any material having similar optical properties as the material of the optical filter 244, including, but not limited to, thickness and refractive index. In one aspect, the transparent material within the first detection well 908 can be fused silica glass of the same thickness as the optical filter 244.

[0128] As a non-limiting example, in Figure 3 The transmission spectra of the OG 530 filter are provided in . Figure 3 As shown, the transmission spectrum of the OG 530 filter overlaps with the emission spectrum of the MB-102 exogenous phosphor and the emission spectrum (emission wavelength) of the green LED used as the second light source 220. In addition, the transmission spectrum of the OG 530 filter does not include the emission spectrum of the blue LED used as the first light source 218 and the absorption spectrum (excitation wavelength) of the MB-102 exogenous phosphor.

[0129] In one aspect, transparent materials such as glass 246 and optical filter 244 can be secured to flanges formed within first detection well 908 and second detection well 910, respectively. Transparent materials such as glass 246 and optical filter 244 can be secured in place using an opaque and / or light-absorbing adhesive, including but not limited to black epoxy, to ensure that all light received through first detector aperture 1004 and second detector aperture 1006 passes through optical filter 244 or glass 246 before being detected by first and second light detectors 222 / 224. In another aspect, the sides of filter 244 or glass 246 can be painted black with a light-absorbing coating, including but not limited to India ink, to ensure that light does not reach first and second light detectors 222 / 224 without passing through optical filter 244 or glass 246.

[0130] In one aspect, due to the Lambertian distribution of the angles of light exiting the patient's skin, the height of the detector wells 908 / 910, combined with the diameter of the detector apertures 1004 / 1006, can limit the fraction of light emitted from the second and third regions 208, 210 of the patient's skin that reaches the active area of the light detectors 222 / 224. In one aspect, the fraction of light emitted from the second and third regions 208, 210 of the patient's skin that is received by the light detectors 222 / 224 can be in the range of about 5% to about 90%. In various other aspects, the fraction of light can be in the range of about 5% to about 15%, about 10% to about 20%, about 15% to about 25%, about 20% to about 30%, about 25% to about 35%, about 30% to about 40%, about 35% to about 45%, about 40% to about 60%, about 50% to about 70%, and about 60% to about 90%.

[0131] In one aspect, for Figure 6 and Figure 7 In the illustrated sensor head 204 having apertures 1002 / 1004 / 1006 having a diameter of 1 mm, approximately 10% of the light emitted from the patient's skin surface can reach the active area of the light detectors 222 / 224 to be detected. In various aspects, the sensor head 204 can further include additional optical elements, including but not limited to lenses and / or prisms, which are configured to compensate for the Lambertian distribution of light angles in order to enhance the proportion of light emitted from the patient's skin that is directed to the active area of the light detectors 222 / 224.

[0132] iii) Temperature sensor

[0133] Reference Figure 2The sensor head 204 may also include one or more additional temperature sensors 228 configured to monitor the temperature of various areas within and near the sensor head 204. Non-limiting examples of suitable areas whose temperatures may be monitored by the one or more additional temperature sensors 228 include: the temperature at the skin surface of the patient 202; the temperature near the first light source 218 and / or the second light source 220; the ambient temperature outside the sensor head 204; the temperature of the housing 600 of the sensor head 204; and any other suitable areas. In one aspect, the additional temperature sensors 228 may be configured to monitor the temperature near temperature-sensitive electrical components, including but not limited to: the light sources 218 / 220, such as LEDs; the light detectors 222 / 224, such as silicon photomultipliers (SPMs); and any other temperature-sensitive electrical components of the sensor head 204. In some aspects, the one or more temperatures measured by the one or more additional temperature sensors 228 may be used as feedback in control methods for one or more temperature-sensitive devices of the system 200, as described below.

[0134] As a non-limiting example, temperature measurements may be used to control the amount of light energy generated by an LED used as the first or second light source 218 / 220. In this example, the temperature sensor 1108 (see FIG. 110 ) may be used in the control scheme. Figure 11 ) to modulate the amount of power supplied to the LED light source to compensate for the effect of LED temperature on the LED light output. In another aspect, an additional temperature sensor 228 can monitor the temperature of the LED light source 218 / 220 to monitor and / or compensate for temperature variations of the LED, as well as to monitor and / or compensate for the temperature-dependent transmission of the optical filter to maintain a relatively constant output wavelength.

[0135] As another non-limiting example, the additional temperature sensor 228 may be in the form of the temperature sensor 816 (see Figure 8 ) is included in the sensor head 204, and the temperature sensor 816 is configured to monitor the temperature of the housing 600 near the contact surface 606 of the sensor head 204. Figure 7 、 Figure 8 and Figure 9 In one aspect, the temperature sensor 816 can be epoxied into the temperature sensor opening 706 in the aperture plate 702. In this aspect, the space 918 between the circuit board (not shown) and the lower housing 604 can be filled with thermal putty to ensure good thermal conduction and heat dissipation.

[0136] In this example, the measured housing temperature can be used to modulate the light output of the sensor head 204 to prevent overheating of the skin of the patient 202 during use. In another aspect, an additional temperature sensor 228 can monitor the temperature of the LED light sources 218 / 220 to monitor and / or compensate for temperature variations of the LEDs, thereby enabling the LED light sources 218 / 220 to maintain a relatively constant output wavelength.

[0137] In another aspect, if an over-temperature condition is detected, the temperature measured by the one or more additional temperature sensors 228 can ensure subject safety by disabling one or more electrical devices including the light sources 218 / 220 and / or the light detectors 222 / 224. In one aspect, an over-temperature condition can be indicated if the housing temperature detected by the temperature sensor 816 is greater than approximately 40°C. In various other aspects, an over-temperature condition can be detected if the housing temperature is greater than approximately 40.5°C or greater than approximately 41.0°C.

[0138] B. Controller

[0139] Refer again Figure 2 In various aspects, the system 200 may include a controller 212 configured to operate the light source 218 / 200 and the light detector 222 / 224 in a coordinated manner to obtain multiple measurements of fluorescence of an exogenous fluorescing agent within tissue of the patient 202, to correct the fluorescence data to remove the effects of dynamic variations in background signals as described below, and to convert the fluorescence measurements into parameters representative of renal function of the patient 202. Figure 11 is a schematic diagram of an electronic circuit 1100 that, in one aspect, illustrates the arrangement of various electrical components capable of operating the system 200. In one aspect, the controller 212 may be a computing device that further includes an operating unit 214 and a display unit 216.

[0140] i) Light source control unit

[0141] Refer again Figure 2 , the controller 212 may include a light source control unit 230 configured to operate the first light source 218 and the second light source 220 to generate light of an excitation wavelength and an emission wavelength, respectively, in a coordinated manner to generate a repetitive pulse sequence, such as Figure 5Schematically illustrated. In various aspects, the light source control unit 230 can generate a plurality of light control signals that encode one or more light control parameters, including, but not limited to: activation or deactivation of each light source; the relative timing of activation and deactivation of each light source to enable light pulse width, pulse repetition rate, electrical power delivered to the light source, or other parameters associated with light pulse flux or light pulse power; other light source-specific parameters that control the light output of the light source; and any other relevant light control parameters. In one aspect, the light source control unit 230 can receive one or more feedback measurements that are used to modulate the plurality of control signals to compensate for variations in the performance of the light source, thereby maintaining a relatively stable output of light from the light source. Non-limiting examples of feedback measurements used by the light source control unit 230 include: light output of the light sources 218 / 220 measured within the source well 902 by the first monitor photodiode 904 and the second monitor photodiode 906, respectively; the temperature of the light sources 218 / 220; and any other feedback measurements relevant to monitoring the performance of the light sources 218 / 220.

[0142] As a non-limiting example, the light source control unit 230 can be configured to operate the LED light sources 218 / 220. In this example, the light output of the LED light sources 218 / 220 can be controlled by controlling the current amplitude provided to each LED. In one aspect, the light source control unit 230 can include at least one waveform generator 1122, which includes but is not limited to a field programmable gate array FPGA having a 16-bit DAC 1124, which is operably coupled to the LED current source 1126, such as Figure 11 In one aspect, a waveform generated by at least one waveform generator 1122, including but not limited to a square wave, can control the output from the LED current source 1126. In one aspect, the current amplitude provided to the LED light sources 218 / 220 can be adjusted based on the waveform signal provided by the waveform generator / FPGA 1122.

[0143] Reference Figure 5 In one aspect, each light pulse train 500 includes an emission wavelength light pulse 502 and an excitation wavelength light pulse 504, both of which are composed of a plurality of square waves 506 generated by the first and second LED light sources 218 / 220. Figure 11 , the square wave generated by waveform generator 1122 is received by LED current source 1126. The current generated by the LED current source includes a square wave similar to the waveform generated by waveform generator 1122. Without being limited to any particular theory, since the intensity of the light generated by the LED light source 218 / 220 is proportional to the amplitude of the received current, the light generated by the LED light source 218 / 220 also includes the following: Figure 5In another aspect, discussed in detail below, the square wave generated by waveform generator 1122 can also be used by acquisition unit 234 in a synchronized detection approach to reduce the effects of various confounding factors, including but not limited to detection of ambient light by detector signals generated by light detectors 222 / 224 during illumination of patient tissue at the emission wavelength and excitation wavelength by first light source 218 / 220 and second light source 218 / 220, respectively.

[0144] In various other aspects, various alternating LED pulse modulation schemes may be equivalently employed without limitation. In one aspect, excitation pulses and emission pulses are delivered in an alternating series, with a dark period separating each pulse. In another aspect, the first and second LED light sources 218 / 220 are each modulated at a 50% duty cycle but at different modulation frequencies, thereby allowing the signals associated with the excitation pulses and emission pulses to be separated by frequency filtering.

[0145] Without being limited to any particular theory, the total optical power delivered to the patient's skin can be limited by at least two factors: photobleaching of exogenous fluorescers and / or endogenous chromophores, and overheating of the patient's tissue irradiated by the system 200. In one aspect, tissue heating can impose an absolute limit of approximately 9 mW on the optical power that can be delivered to the skin, based on safety standards including, but not limited to, ANSI / IESNA RP-27.1-05. In another aspect, photobleaching of skin autofluorescence associated with endogenous chromophores (including, but not limited to, collagen, hemoglobin, and melanin) can provide a background signal to the measured fluorescence that remains relatively constant as long as chromophore autobleaching does not occur. This constant autofluorescence background can be subtracted from the raw fluorescence signal, but if the autofluorescence varies over time due to photobleaching, this background correction may interfere with the kinetic calculation of the renal decay time constant (RDTC). In one aspect, the optical output power of the first light source 218 and / or the second light source 220 can be limited to a level below a power threshold associated with chromophore photobleaching.

[0146] Refer again Figure 9 In various aspects, the light output of the light sources 218 / 220 can be measured using monitor photodiodes 904 / 906. Because the light intensity reaching these monitor photodiodes 904 / 906 is typically much stronger than the light intensity reaching the light detectors 222 / 224 through the patient's skin, less sensitive light detection devices, including but not limited to PIN photodiodes, can be used to monitor the output of the light sources 218 / 220.

[0147] In various aspects, the system 200 can be configured to operate within a range of skin tones observed in the human population. Without being limited to any particular theory, variations in skin tone between different patients 202 can result in variations in the detected fluorescence signal over a range of approximately three orders of magnitude. Additionally, changes in the concentration of the exogenous fluorescent agent within each patient 202 can vary over a range of approximately two orders of magnitude due to elimination of the agent in the kidney over time. In various aspects, the system 200 can be configured to detect fluorescence from endogenous fluorescent agents over an intensity range of more than five orders of magnitude. In these various aspects, the system 200 can be configured by modulating at least one operating parameter, including but not limited to: the amplitude of the light output of the light source 218 / 220 and the sensitivity of the light detector 222 / 224 corresponding to the detector gain.

[0148] In one aspect, the intensity of the light output by the light source 218 / 220 can be manually set by a user via the operating unit 214. In another aspect, the light source control unit 230 can be configured to automatically modulate the intensity of the light generated by the light source 218 / 220. In one aspect, the light source control unit 230 can be configured to control the intensity of the light generated by the LED light source 218 / 220 within a range of normalized output intensities from 0 (off) to 1 (maximum power). In one aspect, the intensity of the light source 218 / 220 can be set by the light source control unit 230 in coordination with the detector gain of the light detector 222 / 224 set by the light detector control unit 232, as described below.

[0149] In one aspect, the light source control unit 230 can use signals obtained during the first 10 detection cycles acquired by the system 200 after data acquisition is initialized but before the injection of an exogenous fluorescent agent to automatically adjust the light intensity generated by the LED light sources 218 / 220 and the gain of the light detectors 222 / 224. In this example, the initial detection cycles can be obtained with the LED light sources 218 / 220 set to approximately 10% of the maximum LED intensity (corresponding to a normalized output intensity of 0.1) and the light detectors 222 / 224 set to a low gain. Based on the detected intensity of light received by the light detectors 222 / 224 at the excitation and emission wavelengths during one detection cycle, the corresponding LED intensity can be modulated so that the analog signal generated by the light detectors 222 / 224 corresponds to approximately ¼ of the full range of each detector's analog-to-digital converter (ADC) at the low detector gain setting. If the signal generated by the light detectors 222 / 224 is inconsistent in response to light generated by the second LED light source 220 at the emission wavelength, a larger signal can be used to modulate the power setting of the second LED light source 220. If the above method results in modulation to an LED intensity setting above the maximum intensity (corresponding to a normalized output intensity of 0.1), the LED intensity setting is set to the maximum setting. Without being limited to any particular theory, the target level of the signal generated by the light detector 222 / 224 (i.e., 1 / 4 of the ADC range) is selected to retain additional light detection capability to detect signals generated by changes in the optical properties of the tissue of the patient 202 during the study due to any one or more of a number of factors (including, but not limited to, the introduction of exogenous fluorescent agents into the patient 202).

[0150] In one aspect described above, once the light source control unit 230 sets the LED intensity in coordination with the detector gain of the light detectors 222 / 224 set by the light detector control unit 232 during the first 10 detection cycles, an additional 10 detection cycles are obtained to confirm the suitability of these settings for the operation of the system 200 given the tissue properties of the particular patient 202, and the LED intensity settings and detector gain are then recalculated as described herein. If the newly calculated LED intensity is within two times the previously determined setting and the detector gain does not change, the previously determined setting is retained for subsequent data acquisition cycles used to determine renal function. Otherwise, the setting is updated using the same method described herein, and an additional 10 data acquisition cycles are performed to confirm the stability of the setting. This process is repeated until either setting is determined to be acceptably stable or 10 data acquisition cycles are performed to obtain the setting, in which case the most recently determined setting is used for all subsequent data acquisitions, and the user can be notified via the display unit 216 that the setting may not be optimal.

[0151] ii) Photodetector control unit

[0152] Refer again Figure 2 , the controller 212 may include a light detector control unit 232 configured to operate the first light detector 222 and the second light detector 224 to be capable of detecting light at the emission wavelength and unfiltered light at all wavelengths, respectively. In various aspects, the light detector control unit 232 may generate a plurality of detector control signals encoding one or more detector control parameters including, but not limited to, detector gain. In various other aspects, the light detector control unit 232 may generate a plurality of light measurement signals encoding the intensity of light detected by the light detectors 222 / 224, including, but not limited to, the intensity of light detected by the light detectors 222 / 224, which in various aspects may be converted by an analog-to-digital converter (ADC) 1102 (see Figure 11 ) received raw detector signal. On the other hand, when the system 200 is configured in engineering mode, the detector gain and / or other detector control signals can be manually set by the user.

[0153] In various other aspects, the amount of light received by the light detectors 222 / 224 can vary due to any one or more of at least several factors, including but not limited to: variations in skin tone observed between individual patients 202, variations in exogenous fluorescer concentration within each patient 202, and any other relevant parameters. In one aspect, the gain of the first light detector 222 and the second light detector 224 can be set by a user via the operating unit 214. In another aspect, the light detector control unit 232 can be configured to generate a bias voltage via the bias voltage generator 1112 (see FIG. Figure 11 )'s bias voltage gain automatically modulates the gain of the photodetectors 222 / 224.

[0154] In one aspect, signals obtained during the first 10 detection cycles obtained by the system 200 after data acquisition is initialized but before injection of an exogenous fluorescent agent are used by the photodetector control unit 232 to automatically adjust the gain of the photodetectors 222 / 224, as well as the output intensity of the light sources 218 / 220. As previously described herein, the initial detection cycles can be obtained with the LED light sources 218 / 220 set to approximately 10% of the maximum LED intensity (corresponding to a normalized output intensity of 0.1) and the photodetectors 222 / 224 having a low gain setting, and the LED intensity can be modulated so that the analog signal generated by the photodetectors 222 / 224 corresponds to approximately ¼ of the full range of each detector analog-to-digital converter (ADC) at the low detector gain setting.

[0155] In this aspect, if the intensity of the first LED light source 218 (generating light at the excitation wavelength) is set to the maximum of the LED power range, a high detector gain for the second light detector 224 corresponding only to filtered measurements at the excitation wavelength can be considered. In various aspects, for a given light detector, a high detector gain can be 10 times higher than the corresponding low detector gain. Without being limited to any particular theory, assuming that the exogenous fluorescent agent is MB-102 introduced into the patient 202 at a dose level of approximately 4 μmol / kg of the patient's body weight, the expected peak detected fluorescence signal from the exogenous fluorescent agent during injection and renal elimination is generally expected to be approximately 10% of the magnitude of the signal received during illumination by the first light source 218 at the excitation wavelength. In one aspect, if the expected detector signal received during maximum LED intensity illumination and with the detector gain set to a high setting is still less than 10% of the range of the detector ADC, the detector gain for that measurement is increased by a factor of 10. In another aspect, the saturation state can persist for a predefined period of time, including but not limited to a 30-second period before adjustments are made to the detector gain or LED power to avoid reacting to stray signal spikes.

[0156] On the other hand, if the detected light signal from one of the photodetectors 222 / 224 exceeds a threshold percentage of the maximum ADC range, the photodetector control unit 232 can adjust the detector gain to a lower gain level to avoid signal saturation. While the highest threshold percentage of the maximum ADC range associated with signal saturation is 100%, the onset of severe detector nonlinearity occurs at a threshold percentage of approximately 40% or higher, while the onset of mild detector nonlinearity occurs when the threshold percentage exceeds approximately 15%. In various aspects, the threshold percentage of the maximum ADC range can be 40%, 35%, 30%, 25%, 20%, 18%, 17%, 16%, 15%, 14%, 13%, 12%, 11%, 10%, 9%, 8%, 7%, 6%, or 5% of the maximum ADC range. In one aspect, if the detected light signal from one of the photodetectors 222 / 224 exceeds approximately 8% of the maximum ADC range, the gain setting is adjusted. As a non-limiting example, if the detector gain is high on a signal near saturation, it is adjusted to low. If the current detector gain is set low and the corresponding detected light signal remains above a threshold percentage of the maximum ADC range, the LED output power setting of the corresponding LED light source can be reduced by a factor of 10.

[0157] In one aspect, the photodetector control unit 232 can receive one or more feedback measurements used to modulate the plurality of detector signals to compensate for variations in photodetector performance due to variations in temperature and / or light source output. Non-limiting examples of feedback measurements used by the photodetector control unit 232 include: light output of the light sources 218 / 220 as measured within the source well 902 by the first monitor photodiode 904 and the second monitor photodiode 906, respectively (see FIG. Figure 11 ), the temperature of the light detector 222 / 224 measured by the first temperature sensor 1106, the LED temperature measured by the second temperature sensor 1108, the temperature of the sensor head housing measured by the third temperature sensor 1128, the LED supply current from the LED current source 1126, and any other feedback measurements relevant to monitoring the performance of the light detector 222 / 224.

[0158] In various aspects, the light detectors 222 / 224 can be silicon photon multiplier (SPM) detectors, which can include low-noise internal amplification and can operate at lower light levels relative to other light sensor devices such as PIN photodiodes. The detector signals generated by the SPM detectors 222 / 224 can be amplified using transimpedance amplifiers 1120 / 1118, respectively (see FIG. Figure 11 ) to convert the current generated by each SPM photodetector 222 / 224 into a measurable detector voltage. The transimpedance amplifier 1118 on the second SPM photodetector 224 (i.e., detecting only filtered light at the excitation wavelength) can include a switchable detector gain that can select a low gain configured to detect a larger dynamic range for fluorescence measurement when the first LED light source 218 is activated to generate light at the emission wavelength. When the second light source 220 is not activated, the switchable detector gain can further select a high gain setting for the second SPM photodetector 224 to enhance the sensitivity of the second SPM photodetector 224 during the phase of the detection cycle when detecting light at the emission wavelength generated by exogenous fluorescent agents within the tissue of the patient 202 to ensure that the expected dark current from the second SPM photodetector 224 occupies less than 1 / 4 of the total ADC output range. In one aspect, the second transimpedance amplifier of the second SPM photodetector 224 may include a low detector gain configured to provide a transimpedance gain of approximately 4 kΩ, corresponding to approximately twice the value of the transimpedance resistor due to differential operation, and may also include a high detector gain configured to provide a transimpedance gain of approximately 40 kΩ. In another aspect, the first transimpedance amplifier of the first SPM photodetector 222 may include a fixed detector gain configured to provide a transimpedance gain of approximately 2 kΩ.

[0159] iii) Collection unit

[0160] Refer again Figure 2 In various aspects, the controller 212 may also include an acquisition unit 234. The acquisition unit 234 may be configured to receive a plurality of signals from the light sources 218 / 220, the light detectors 222 / 224, the additional light detector 226, and the additional temperature sensor 228, and process the plurality of signals to generate one or more raw signals, including, but not limited to, a raw fluorescence signal encoding the fluorescence intensity detected by the second light detector 224 during illumination at the excitation wavelength, and a raw internal reflection signal corresponding to the light intensity at the excitation wavelength detected by the first light detector 222 during illumination at the excitation wavelength, and the light intensity at the emission wavelength detected by both light detectors 222 / 224 during illumination at the emission wavelength.

[0161] The plurality of signals received from the various sensors and devices described above are typically analog signals including, but not limited to, voltage and current. In various aspects, the acquisition unit 234 may be capable of sending the analog signals to one or more analog-to-digital converters (ADCs) to convert the analog signals into digital signals for subsequent processing by the processing unit 236. Figure 11 is a schematic diagram of a circuit 1100 illustrating the arrangement of various electrical devices and components of the sensor head 204. In one aspect, an analog signal encoding the intensity of light detected by the first and second light detectors 222 and 224 may be received by the first ADC 1102.

[0162] In various aspects, at least one 24-bit Σ-Δ ADC can be used to digitize the analog signals generated by the photodetectors 222 / 224 and the various monitor sensors. Figure 11In one aspect, a high-speed 24-bit Σ-Δ ADC 1102 can be used to digitize analog signals encoding measurements from time-sensitive sensors. In this aspect, time-sensitive sensors include sensors associated with the generation and detection of light pulses characterized by potentially rapidly varying signals. Non-limiting examples of time-sensitive sensors of system 200 include: first and second light detectors 1118 / 1120, and first and second monitor photodiodes 904 / 906. In another aspect, a low-speed 24-bit Σ-Δ ADC 1104 can be used to digitize analog signals encoding measurements from less time-sensitive sensors. In this other aspect, less time-sensitive sensors include sensors associated with monitoring system conditions characterized by typically slow-varying signals, including, but not limited to, the temperature of various system components and / or areas. Non-limiting examples of less time-sensitive sensors of system 200 include first and second temperature sensors 1106 / 1108 configured to monitor the temperatures of light detectors 222 / 224 and light sources 218 / 220 , respectively, and a third temperature sensor 1128 configured to monitor the temperature of housing 600 of sensor head 204 .

[0163] In various aspects, the acquisition unit 234 can also be configured to enable synchronous detection of light by the detectors 222 / 224. Without being limited to any particular theory, the synchronous detection method is believed to suppress noise from the detector signal associated with the detection of light generated by the light sources 118 / 120 and fluorescence generated by exogenous fluorescing agents within the tissue of the patient 202 by distinguishing the detector signal from noise associated with detecting ambient light or other interfering sources.

[0164] Figure 12 is a schematic diagram of a synchronous detection method in one aspect. Figure 11 and Figure 12 The waveform generator / FPA 1122 can generate a digital square wave 1202 that is received by the DAC 1124, and the resulting analog-converted square wave is received by the LED current source 1126. The resulting current generated by the LED current source 1126, also characterized by a waveform proportional to the analog-converted square wave, drives the LED light sources 218 / 220. The light generated by the LED light sources 218 / 220, after passing through the tissue of the patient 202, is detected by the photodetectors 222 / 224 along with the fluorescence generated by the endogenous fluorescer, and is digitized by the high-speed ADC 1102.

[0165] Refer again Figure 11 and Figure 12 The digital square wave 1202 generated by the waveform generator / FPA 1122 can also be generated by the DAC 1110 (see Figure 11) is converted into an in-phase reference sine wave 1210 and an out-of-phase / quadrature reference cosine wave 1212. In one aspect, the digitized detector signal from the ADC 1102 and the in-phase reference sine wave 1210 can be sampled and sign-multiplied at a first multiplier 1214 to generate a plurality of in-phase modulated signals. Additionally, the digitized detector signal and the quadrature reference cosine wave 1212 can be sampled and sign-multiplied at a second multiplier 1216 to generate a plurality of quadrature (out-of-phase) modulated signals. In this regard, the acquisition unit 234 can delay the samples from the reference waves 1210 / 1214 by an amount equal to the relative delay between the DAC 1124 generating the reference waves 1210 / 1214 and the ADC 1102 digitizing the detector signal to synchronize the reference waves 1210 / 1214 with the detector data being acquired.

[0166] Refer again Figure 12 , the in-phase modulated signals may be summed in a first accumulator 1218 to generate an in-phase intensity signal 1224. Similarly, the quadrature modulated signals may be summed in a third accumulator 1222 to generate a quadrature intensity signal 1228. The raw digitized detector signals may also be summed in a second accumulator 1220 to generate an average intensity signal 1226. Furthermore, the in-phase intensity signal 1224 and the quadrature intensity signal 1228 may be square root summed to generate an amplitude signal 1230.

[0167] Without being limited to any particular theory, the integration interval of the accumulators 1218 / 1220 / 1222 can correspond to an integer number of modulation periods (corresponding to the period of the digital square wave 1202) to avoid biasing the measured signal. The phase accumulators 1218 / 1220 / 1222 used to control synchronous detection operate at integers, but the sampling clock frequency and the modulation frequency are not integer-divisible, so the number of periods is not an exact integer. However, the error associated with this mismatch can be minimized by adjusting the actual modulation frequency to match the achievable sampling interval as closely as possible and allocating an appropriate number of bits to the phase accumulator. In one aspect, the error associated with the mismatch between the modulation frequency and the sampling interval can be as small as 10 6 On the order of about a portion of the .

[0168] In one aspect, the digital square wave 1202 used to modulate the LED light sources 218 / 220 and implement the synchronous detection method described herein is generated at a frequency of approximately 1 kHz. Without being limited to any particular theory, the square wave is selected as the modulating waveform to enhance the signal-to-noise ratio (SNR) compared to a pure sine wave as the modulating waveform at the same peak power level.

[0169] In another aspect, acquisition unit 234 can also be configured to demodulate in-phase intensity signal 1224, average intensity signal 1226, and quadrature intensity signal 1228. In one aspect, acquisition unit 234 can single out each component at a fundamental frequency characterized by an amplitude that is (4 / π) times the amplitude of square wave 1202 used to modulate intensity signals 1224 / 1226 / 1228. In various aspects, to suppress 50 / 60 Hz electrical noise generated by the AC power supplies, and corresponding 100 / 120 Hz optical noise generated by ambient light sources powered by these power supplies, the integration period of accumulators 1218 / 1220 / 1222 can be selected to be a multiple of 100 ms. In these various aspects, the selected integration period ensures that integration of accumulators 1218 / 1220 / 1222 occurs over an integer number of cycles of the 50, 60, 100, and 120 Hz signals.

[0170] iv) Processing Unit

[0171] Refer again Figure 2 In various aspects, the controller 212 may also include a processing unit 236 configured to apply corrections to the demodulated detector signal and convert selected portions of the corrected detector signal into a measure of renal function. Figure 13 is a block diagram illustrating, in one aspect, subunits of the processing unit 236. Figure 13 , the processing unit 236 may include a pre-processing sub-unit 1302 configured to determine and correct the detector signal to remove signal artifacts associated with various confounding effects, including but not limited to physiologically induced signal variations, variations in the power supplied to the light sources 218 / 220, nonlinearities in the detector response, ambient temperature variations, and tissue heterogeneity. The processing unit 236 may also include a background subtraction sub-unit 1304 configured to eliminate a portion of the detector signal attributable to background factors, such as autofluorescence of the tissue and / or leakage of light at the excitation wavelength through the optical filter 244 of the second light detector 224. The processing unit 236 may additionally include a background correction sub-unit 1306 configured to apply a background correction method to eliminate the effects of dynamic variations in the background signal associated with autofluorescence and / or leakage of excitation wavelength light to the second light detector 224 configured to detect only emission wavelength light, and apply the background correction to the first detector, DR ex,meas Transition to DR ex,photonsThe processing unit 236 may further include a post-reagent administration selection subunit 1308 configured to select a portion of the detector data associated with the post-equilibration period for subsequent analysis to determine the patient's renal function. The processing unit 236 may further include an RDTC calculation subunit 1310 configured to convert the detector signal obtained during the post-equilibration period to generate a renal decay time constant indicative of the patient's renal function. The processing unit 236 may further include a fault detection subunit 1312 configured to monitor the amplitude of the detector signal to detect any faults in the system.

[0172] -Preprocessing subunit

[0173] In one aspect, various modules of the pre-processing subunit 1302 are used to pre-process the raw signal corresponding to the light intensity detected by the light detector 222 / 224, where the light intensity detected by the light detector 222 / 224 corresponds to the illumination of the first light source 218 and the second light source 220 at the excitation wavelength and the emission wavelength, respectively, to eliminate the effects of multiple confounding factors from the raw signal, thereby producing a signal that more accurately reflects the potential specific signal of interest.

[0174] As a number of non-limiting examples, the intensity of light generated by a light source can vary due to one or more of a number of factors, including, but not limited to, fluctuations in the current supplied to the light source and changes in the ambient temperature of the light source. Light represented by two or more wavelengths emitted by the same source aperture of a sensor head may not share the same path to the same detector. Detectors may have thermally dependent sensitivity and gain. Furthermore, the optical filter associated with the second light detector 224 may have temperature-dependent transmission characteristics.

[0175] In one aspect, the pre-processing subunit 1302 is configured to process the raw signals corresponding to the light intensities detected by the first and second light detectors 222 / 224 to remove one or more of the measurement errors associated with the devices and elements of the system 200 and patient-specific factors, including but not limited to the plurality of factors described above. Figure 22A is a block diagram illustrating modules of the pre-processing subunit 1302 in one aspect. Figure 22B is a block diagram illustrating modules of the pre-processing sub-unit 1302a in the second aspect.

[0176] In one aspect, as Figure 22AAs shown, the pre-processing subunit 1302 1) resamples the signal using the method of the resampling module 2202 described below, 2) eliminates the saturated detector signal using the method of the detector output saturation detection and elimination module 2204 described below, 3) corrects the temperature-dependent detector gain using the method of the detector temperature correction module 2206 described below, 4) corrects the signal for instrument light directionality using the method of the light directionality correction module 2208 described below, 5) corrects the signal for filter throughput and temperature-dependent variations in fluorescence using the method of the filter throughput temperature correction (emission) module 2212 described below, 6) corrects tissue non-uniformity using the method of the tissue non-uniformity correction module 2216 described below, 7) corrects the signal and signal decomposition for filter throughput and temperature-dependent variations in excitation light using the method of the filter throughput temperature correction (excitation) module and signal decomposition module 2214 described below, and 8) corrects optical power variations using the method of the fractional photon normalization module 2218 described below.

[0177] In one aspect, as Figure 22B As shown, the pre-processing subunit 1302a uses the method of the detector temperature correction module 2206a described below to calculate the signal magnitude, uses the method of the resampling module 2202a described below to resample the signal, uses the method of the detector output saturation detection and elimination module 2204a described below to eliminate saturated samples, uses the method of the detector temperature correction module 2206a described below to correct the signal for temperature-dependent detector gain, uses the method of the partial photon normalization module 2218a described below to correct the signal for optical power variation, uses the filter throughput temperature correction (excitation) module and signal decomposition module 2214a described below to correct the leakage of excitation light onto the measured fluorescence signal, and uses the filter throughput temperature correction (emission) module 2212a described below to correct the leakage of fluorescence onto the measured excitation diffuse reflectance signal.

[0178] -Resampling module

[0179] Reference Figure 22A and Figure 22B, the pre-processing subunit 1302 / 1302a includes, in various aspects, a resampling module 2202 / 2202a, which is configured to reduce signal variations associated with physiological processes of the patient 202, including but not limited to heartbeat and respiration. Typically, an acquisition sequence is characterized by alternating illumination intervals at excitation and emission separated by no illumination intervals (i.e., dark intervals). Although the two illumination intervals (excitation / emission) are timestamped with the same timestamp value as described above, the dark interval between the excitation and emission illumination intervals results in a separation interval between the excitation and emission illumination intervals. Without being limited to any particular theory, if the separation interval associated with the acquisition sequence is on the order of the separation interval between physiological events (such as heartbeat or respiration), physiological noise can be introduced into the signal. In various aspects, this physiological noise can be reduced by resampling the signals associated with the excitation and emission illuminations to overlap prior to subsequent processing of the signals.

[0180] As a non-limiting example, a sample sequence may include a 100 ms dark interval, a 100 ms illumination interval at the excitation wavelength, a second 100 ms dark interval, and a 100 ms illumination interval at the emission wavelength. Each sample data packet is recorded with a single timestamp, and each sample data packet is separated by a 400 ms interval. Because physiological signal changes, such as those from a heartbeat, occur on the same time scale, a 200 ms difference between the acquisition of signals associated with the excitation wavelength and the emission wavelength becomes apparent in the signal. A pre-processing subunit 1302 may be used to reduce physiological signal noise by first resampling the signals associated with illumination at the excitation wavelength and the emission wavelength to overlap before performing any additional signal processing as described below. In this non-limiting example, the signal associated with illumination at the excitation wavelength may be shifted forward by 100 ms, and the signal associated with illumination at the emission wavelength may be shifted backward by 100 ms, resulting in an overlap of the signals.

[0181] In various aspects, the resampling module 2202 performs resampling as described above on the signals detected by both the first and second detectors 222 / 224. In one aspect, the resampling module 2202 acts as a form of low pass filter.

[0182] -Detector output saturation detection and elimination module

[0183] Refer again Figure 22A and Figure 22BIn various aspects, the pre-processing subunit 1302 / 1302a includes a detector output saturation detection and elimination module 2204 / 2204a configured to detect and eliminate signal values that fall outside the detection range of the light detector 222 / 224. In one aspect, the pre-processing subunit 1302 compares the detected signals to the maximum ADC signal. If any signal falls within a threshold range of the maximum ADC signal using an average or peak signal value, the detector output saturation detection and elimination module 2204 identifies the value and eliminates it from further processing.

[0184] -Detector temperature correction module

[0185] Refer again Figure 22A and Figure 22B In various aspects, the pre-processing subunit 1302 / 1302a includes a detector temperature correction module 2206 / 2206a configured to perform temperature correction to compensate for the thermal sensitivity of the photodetectors 222 / 224. In one aspect, the intrinsic detector gain of a silicon photomultiplier (SPM) device, typically used as a photodetector, is coupled to the device breakdown voltage and bias voltage generator 1112 (see Figure 11 ) is proportional to the difference between the applied bias voltage (referred to herein as overvoltage). In this regard, the breakdown voltage varies with temperature in a well-characterized manner. In one aspect, temperature correction accounts for both this internal detector gain variation and additional temperature-related variations in photon detection efficiency.

[0186] In one aspect, the temperature correction can be a scaling correction applied to the detector measurement, where the scaling correction is based on the measured detector temperature. In one aspect, the measured photodetector signal can be divided by a calculated gain G(t) to remove the temperature dependency. The scaling correction G(t) can be calculated according to equation (2):

[0187]

[0188] In equation (2), the first temperature sensor 1106 (see FIG. 1 ) configured to monitor the temperature of the sensors 222 / 224 is Figure 11 ) to obtain the monitored temperature T. Bias voltage (V bias ) can be measured by the bias voltage generator 1112. The breakdown voltage (V breakdown ) and the reference temperature (T0) are constants specific to the particular photodetector device included in the system 200. As a non-limiting example, if the photodetectors 222 / 224 are silicon photomultiplier (SPM) devices, then V breakdownmay be 24.5 V and T0 may be 21° C. On the other hand, the coefficient C used in equation (2) may be derived empirically based on measurements obtained using a constant body phantom over an ambient temperature range of about 18° C. to about 26° C. v and C T .

[0189] On the other hand, the temperature portion of the gain correction is determined by equations (3) to (5).

[0190]

[0191]

[0192]

[0193] This gain correction may be applied to each signal amplitude measured by the first and second photodetectors 222 / 224 as follows:

[0194]

[0195] In one aspect, the measured amplitude from each detector and monitor photodiode is calculated as the square root sum of the in-phase amplitude signal 1230 (I) and the quadrature amplitude signal 1232 (Q) according to equation (1):

[0196]

[0197] The signal amplitudes from the photodetectors 222 / 224 calculated using equation (1) are normalized by the monitor photodiode amplitudes for each measurement set corresponding to measurements obtained during illumination of one of the LED light sources 218 / 220 at either the excitation wavelength or the emission wavelength. Because the two monitor photodiodes 904 / 906 can be used with two LED light sources 218 / 220 (see Figure 9 ) are located in the same source well 902, so the average of the two monitor photodiode amplitudes from the corresponding measurement group is used.

[0198] In one aspect, the in-phase intensity signal 1224, the quadrature intensity signal 1228, and the average intensity signal 1226 (see Figure 12 ) are further processed for the number of accumulated samples and ADC scaling so that the intensity signals 1224 / 1226 / 1228 are returned as part of the full range of the high speed ADC 1102 (i.e., the range from a minimum value of 0 to a maximum value of 1). Monitor photodiodes 904 / 906 (see Figure 11 ) is similarly scaled to a fraction of the full range of the low speed ADC 1104.

[0199] In one aspect, Gcorrection Power correction can be incorporated to correct for the effects of fluctuations in the LED power supply. In this regard, the signals from the first monitor photodiode 904 and the second monitor photodiode 906 are calibrated by measuring the light output power with a power meter as the light intensity from the light source 218 / 220 varies. source1 and C source2 The calibration factor is calculated for each recorded monitor photodiode signal value in milliwatts measured by the detector. source1 and C source2 Used to determine the absolute light output into tissue at each wavelength.

[0200] Refer again Figure 22B The detector temperature correction module 2206a uses the LED output signal PD measured by the first monitor photodiode 904 and / or the second monitor photodiode 906 to correct the temperature of the LED. magnitude The temperature corrected detection signal is normalized to correct the signal amplitude for the varying intensity of the LED. In this case, the G from each light source 218 / 220 above correction The variables are modified as follows:

[0201]

[0202] -Light Directivity Correction Module

[0203] Refer again Figure 22A , the pre-processing subunit 1302 in this aspect includes a light directionality correction module 2208 that is configured to correct for variations in the detection signal associated with differences in scattering and absorption of light of different wavelengths by the tissue of the patient 202 during data acquisition. In one aspect, the correction term for light directionality can be measured by acquiring data from one or more homogeneous tissue phantoms and using a sensor configuration in which no emission filter is present. The measured ratio of the signal detected by the first light detector 222 (Det1) to the signal detected by the second light detector 224 (Det2) is used to determine a coefficient G of the signal. ex or G em , the coefficient G of the signal ex or G em These coefficients are used to modify the signal detected by the first light detector 222. In one aspect, the first light detector 222 uses the coefficients G ex or G em The correction of the signal collected in the homogeneous medium is such that the signals measured by the first detector 222 / 224 and the second detector 222 / 224 are equivalent to within 20% of each other. In other aspects, the first light detector 222 uses the coefficient Gex or G em The correction of the signal collected in the homogeneous medium causes the signals measured by the first detector 222 / 224 and the second detector 222 / 224 to be equivalent to about 10%, about 5%, about 2%, and about 1%.

[0204] -Detector nonlinear response correction module

[0205] Refer again Figure 22A In this aspect, the pre-processing subunit 1302 includes a detector nonlinear response correction module 2210 configured to correct for variations in the detection signal associated with the nonlinear response of the detector. In this aspect, the amplitude data obtained by the detectors 222 / 224 can be scaled using a calibration curve based on averaged data.

[0206] -Filter Throughput Temperature Correction (TX) module

[0207] Refer again Figure 22A , in this aspect the pre-processing subunit 1302 includes a filter throughput temperature correction (emission) module 2212 configured to correct for variations in the detection signal associated with temperature-dependent optical properties of the optical filter 244 associated with the second light detector 224 during illumination at the emission wavelength. In this regard, the signal Det2 detected by the second light detector 224 may be corrected according to equation (8):

[0208]

[0209] In various aspects, the signal Det2 measured by the second light detector 224 can be monitored while the ambient temperature is cycled within a range that includes the operating temperature range or a sufficiently large subset of the range to fully determine the temperature dependence of the emission filter. These data are acquired from a uniform, non-fluorescent phantom using an optical filter 244 mounted on the second light detector 224. In addition, simultaneous measurements are monitored from the first light detector 222, and the ratio of the measured values Det2 / Det1 is determined. The nominal filter coefficient C emF,nom is calculated at the nominal operating temperature T nom In this regard, the coefficient C is obtained from the slope of Det2 / Det1 obtained in the ambient temperature range during illumination with the emission wavelength of a uniform non-fluorescent phantom. emF,slopeT .

[0210] - Tissue heterogeneity correction module

[0211] Refer again Figure 22AIn this respect, the pre-processing subunit 1302 includes a tissue heterogeneity correction module 2216 configured to correct variations in the detection signal associated with heterogeneity of tissue intervening between the first region 206 illuminated by the light source 218 / 220 and the second and third regions 208 / 210 in which the light detectors 222 / 224 are located. In this respect, the signal Det1 corrected for light directivity by the light directivity correction module 2208 and the signal Det2 corrected for filter effects by the filter throughput temperature correction (emission) module 2212 are used to calculate C according to equation (9). hetero , which is the coefficient correcting for tissue heterogeneity:

[0212] C hetero =Det2 / Det1 Equation (9)

[0213] - Filter throughput temperature correction (excitation) and signal decomposition modules

[0214] Refer again Figure 22A In this aspect, the pre-processing subunit 1302 includes a filter throughput temperature correction (excitation) module and signal decomposition module 2214, which is configured to correct for changes in the detection signal associated with the temperature-dependent optical properties of the optical filter 244 associated with the second light detector 224 during excitation wavelength illumination. In this aspect, because the emission filter is configured to block light at the excitation wavelength, the filter throughput temperature correction (excitation) module and signal decomposition module 2214 performs correction for changes in the amount of excitation light leakage due to temperature-related changes in the optical properties of the optical filter 244. In addition, the filter throughput temperature correction (excitation) module and signal decomposition module 2214 can correct the signal measured by the first light detector 222 during excitation wavelength illumination due to the presence of fluorescence caused by the excitation wavelength illumination superimposed on a portion of the signal associated with the excitation wavelength illumination.

[0215] In this regard, the effect of changes in temperature dependence on the leakage of the excitation wavelength through the optical filter 244 is calculated as shown in equation (10):

[0216] C exLT =C exLT,nom +C exLT,slopeT (TT nom ) Equation (10)

[0217] In this regard, during illumination at the emission wavelength at the nominal operating temperature T nom C is calculated from the ratio of the signals Det1 and Det2 measured from a uniform non-fluorescent phantom. exLT,nom . C exLT,slopeTis calculated as the slope of the signal Det2 measured from a uniform non-fluorescent phantom over the operating temperature range T during illumination with the wavelength.

[0218] In this regard, the filter throughput temperature correction (excitation) module and signal decomposition module 2214 also performs signal extraction to isolate the portion of the detection signal associated with diffuse reflection of the excitation wavelength illumination and fluorescence. ex2 , that is, the amount of excitation light incident on the second light detector 224 in the absence of the optical filter 244 is not measurable due to the presence of the optical filter 244. In addition, the signal Det1 measured by the first light detector 222 is the signal from the excitation wavelength illumination DR ex1 The composite signal of diffuse reflectance and fluorescence Flr1 is obtained using the tissue heterogeneity correction module 2216 as described above. Hetero The underlying signal is extracted using the following system of equations:

[0219] Det2=C exLT DR ex2 +Flr2 Equation (11)

[0220] Det1=DR ex1 +Flr1 Equation (12)

[0221] Flr2=C Hetero Flr1 Equation (13)

[0222] DR ex2 =C Hetero DR ex1 Equation (14)

[0223] In this regard, Flr2 is determined by solving the above system of equations using only the measurable signals Det1 and Det2 as follows:

[0224] Det2=C exLT C Hetero DR ex1 +Flr2 Equation (15)

[0225] Det2=C exLT C Hetero (Det1-Flr1)+Flr2 Equation (16)

[0226] Det2=C exLT C Hetero Det1-C exLT C Hetero Flr1+Flr2 Equation (17)

[0227] Det2-C exLT C HeteroDet1=Flr2(1-C exLT ) Equation (18)

[0228]

[0229] In this regard, once Flr2 is obtained as described above, the other signals Flr1, DR can be easily obtained by substituting into the above equations (Equations (11) to (14)). ex1 and DR ex2 .

[0230] -Partial photon normalization module

[0231] Refer again Figure 22A In this aspect, the pre-processing subunit 1302 includes a partial photon normalization module 2218, which is configured to convert the detector signal into partial photon units after pre-processing as described above for use in subsequent background subtraction and intrinsic fluorescence correction algorithms, as described herein. In this regard, the detector signal can be converted to photocurrent by reversing the scaling associated with the ADC and the transimpedance amplifier used to acquire the detected signal to obtain a signal in units of photocurrent. Once the photocurrent is obtained, the detector responsivity provided by the manufacturer of the photodetector is used to convert the detector photocurrent into units of watts. The detector signal in watts is then proportional to the source power in watts, which is measured by an additional photodetector 226 used to monitor the output of the light source 218 / 220 to obtain the number of detected partial photons.

[0232] -Optical power correction module

[0233] Refer again Figure 22A and Figure 22B In this regard, the pre-processing subunit 1302 / 1302a includes a partial photon normalization module 2218 / 2218a, which is configured to convert the detector signal into partial photon units after pre-processing as described above for use in subsequent background subtraction and intrinsic fluorescence correction algorithms, as described herein. In this regard, the detector signal can be converted into a photocurrent by reversing the scaling associated with the ADC and the transimpedance amplifier used to acquire the detected signal to obtain a signal in units of photocurrent. Once the photocurrent is obtained, the detector responsivity provided by the manufacturer of the photodetector is used to convert the detector photocurrent into units of watts. The detector signal in watts is then proportional to the source power in watts, which is measured by an additional photodetector 226 used to monitor the output of the light source 218 / 220 to obtain the number of detected partial photons.

[0234] -Excitation light leakage subtraction module

[0235] Refer again Figure 22B In this aspect, the pre-processing subunit 1302a includes a partial photon normalization module 2222, which is configured to normalize the Flr meas In order to obtain the signal composed only of fluorescence photons (Flr photons ) generated by the fluorescence signal, excitation leakage subtraction is performed. In order to eliminate the contribution of excitation light, the excitation leakage is regarded as diffuse excitation part of the signal where the universal calibration factor, C ExLT , make sure to start from Flr meas The part of the signal that is subtracted from the original signal is shown below:

[0236]

[0237] Among them C ExLT is a calibration factor obtained by calculating the ratio between the excitation light detected by two detectors on a non-fluorescent optical model as follows:

[0238]

[0239] Then from Flr meas This signal is subtracted from the fluorescence signal to provide the fluorescence signal due only to the fluorescence photons as follows:

[0240] Flr photons =Flr meas -ExLT

[0241] - Fluorescence light leakage subtraction module

[0242] Refer again Figure 22B In this respect, the pre-processing subunit 1302a includes a sub-unit configured to process Flr meas The signal is subjected to fluorescence leakage subtraction by the fluorescence leakage subtraction module 2224a. In order to obtain the diffuse reflectance, which is defined herein as the signal due to the excitation photons (DRex), the fluorescence leakage subtraction module 2224a is used to perform fluorescence leakage subtraction. photons ) caused by the excitation signal, perform fluorescence leakage subtraction. In order to eliminate fluorescence leakage, the amount of fluorescence leakage observed on the human subject data database is compared with the amount of fluorescence leakage obtained by diffuse reflection, emission signal, The calibration factor, C, is determined by looking at the relationship between the measured tissue heterogeneity and the relationship between the FlrLT . The relationship is linear as shown below:

[0243]

[0244] wherein in one aspect pi and p2 are approximately 0.61 and 0.01, respectively, as determined by the above relationship. In another aspect, without being limited to being defined by the above relationship, pi and p2 may assume any other values.

[0245] DRex is then calculated by subtracting this portion of the measured fluorescence from the diffuse reflectance excitation signal photons The signal is as follows:

[0246]

[0247] b) Baseline subtraction subunit

[0248] Refer again Figure 13 , the processing unit 236 also includes a baseline subtraction subunit 1304. In one aspect, the baseline subtraction subunit 1304 subtracts a baseline signal from the light detector measurements to correct for the effects of autofluorescence and light leakage. As used herein, a baseline period refers to an initial time period of measurements obtained prior to injection of an exogenous fluorescent agent. During the baseline period, it can be assumed that the fluorescence signal measured by the system 200 is associated with tissue autofluorescence and / or excitation light from the LED light source 218 / 220 that leaks through the optical filter 244 of the second light detector 224. In one aspect, the average signal measured during the baseline period (referred to herein as the baseline signal) can be subtracted from subsequent fluorescence measurements to produce a measurement associated only with fluorescence produced by the exogenous fluorescent agent within the patient's tissue.

[0249] In another aspect, correction for excitation light leakage and autofluorescence can be implemented in conjunction with background correction sub-unit 1306. In this aspect, background correction sub-unit 1306 can dynamically calculate the effects of excitation light leakage and autofluorescence during each data acquisition cycle, rather than subtracting the average signal measured during the baseline period. As a result, the subtraction of the excitation light leakage effect can be performed before the diffuse reflectance correction described below, and the subtraction of the autofluorescence effect can be updated by the background correction sub-unit 1306 during each data acquisition cycle.

[0250] c) Background correction subunit

[0251] In one aspect, the background correction subunit 1306 can correct the measured fluorescence data to remove the effects of changes in the optical properties (absorption and scattering) of the tissue of the patient 202 during monitoring of renal extraction of exogenous fluorescent agents within the patient's tissue. As described above, the optical properties of the tissue can change due to any one or more factors including, but not limited to, vasodilation, vasoconstriction, oxygen saturation, hydration, edema, and any other suitable factors within the region of interest being monitored by the system that are associated with changes in the concentration of endogenous fluorophores (such as hemoglobin, collagen, and melanin).

[0252] In one aspect, the background correction subunit 1306 can determine intrinsic autofluorescence (IF auto ) signal, which represents the emission wavelength light emitted by endogenous fluorophores within the patient's tissue during data acquisition. In this regard, IF auto Signal from IF bkrnd The mean or median value of the background intrinsic fluorescence data before reagent injection is obtained. bkrnd The signal looks like this:

[0253]

[0254] The coefficients bkx, bkm and kmFilt are obtained by the global error surface method.

[0255] In one aspect, the values of the exponents used in the above equations are empirically determined using a global error surface method. The method of this aspect comprises providing a diffuse reflectance signal (DR ex , DR em , DR em,filtered ) selects a range of values for each index (bkx, bkm, bkmFilt). In various aspects, the range of values for each index may be affected by any one or more of a variety of factors, including, but not limited to: the design of system 200, including the design of sensor head 204; the properties of the selected exogenous fluorescer, such as excitation wavelength / emission wavelength, absorption efficiency, emission efficiency, and initial dose concentration in the patient's tissue; the type of patient 202 and the corresponding concentration of endogenous chromophores; the location of sensor head 204 on patient 202; and any other relevant factors.

[0256] In one aspect, the method can include selecting a wide range for each coefficient (bkx, bkm, bkmFilt) and performing a wide search. The error surface from the wide search can be analyzed to locate wells in the error surface and the associated range for each coefficient. In this aspect, the method includes adjusting the range of each coefficient to include the region from the wide search, observing the wells in the error surface within the wide search, and repeating the analysis. The method can be iterated until a suitably fine resolution is achieved that accurately achieves the minimum error.

[0257] A step size may be selected for the range of values selected for each index (bkx, bkm, bkmFilt) at 1404. In one aspect, the step size for each factor may be selected based on any one or more of at least a plurality of factors, including, but not limited to: the expected sensitivity of the IF values calculated above to changes in each factor; the total number of suitable combinations of exponents for calculating the factor IF under consideration (including available computing resources, acceptable data processing times, or any other relevant factors); and any other suitable criteria for step size.

[0258] In various aspects, the step size can be the same value for all indices (bkx, bkm, bkmFilt). As a non-limiting example, the step size for all indices can be 0.5. In various other aspects, the step size can be constant for all values of a single index (bkx, bkm, bkmFilt), but the step size selected for each index can vary between different indices. As a non-limiting example, the selected step size for bkx can be 0.01, and the selected step size for bkm and bkmFilt can be 0.6. In various additional aspects, the step size within one or more indices can vary within a range of values for each index. In these various additional aspects, the step size can be reduced within a sub-range of values for an index for which the IF calculated above is predicted to be more sensitive to small changes in the index. Non-limiting examples of suitable step sizes that vary within a range of values for a single index include: different step sizes selected by the user, random step sizes, linear increases and / or decreases in step size, a nonlinear distribution of different step sizes, such as a logarithmic distribution of step sizes, an exponential distribution, or any other suitable nonlinear distribution.

[0259] The selected index range and the selected step size can be used to form a vector of potential values for bkx, bkm, and bkmFilt. For each combination of indices between all vectors, IF is obtained from the measured values Flr, DR using the above equation. ex , DR em and DR em,filtered Calculated. For each combination of indices, multiple IF values are calculated, where each IF value corresponds to one data acquisition cycle. As a non-limiting example, using the vector of potential indices listed above, a total of 405 (5×9×9) IF signals will be calculated.

[0260] In one aspect, a combination of multiple potential indices can be evaluated to select a combination of indices from a plurality to allocate for subsequent diffuse reflectance correction calculated using the above equation. An error estimate of the corrected Flr signal data (i.e., the IF signal data calculated using the above equation) can be calculated. Any estimate of the error can be calculated, including but not limited to an amount related to the residual error of the IF signal data relative to a curve fit of the IF signal data. Any type of known curve fitting method can be used to curve fit the IF signal data, including but not limited to single exponential curve fitting. Without being limited to any particular theory, it is believed that the clearance rate of exogenous fluorescent agents (such as MB-102) from the kidney is expected to be constant, with a renal decay time constant RDTC as the characteristic exponential decay.

[0261] Then intrinsic autofluorescence (IF auto ) is only IF bkrnd The autofluorescence signal Flr is then projected by performing an inverse background diffuse reflectance correction auto , as shown below:

[0262]

[0263] The autofluorescence signal Flr is then eliminated from the measured fluorescence signal Flr auto To determine the specific fluorescence of the reagent (IF) representing the emission wavelength emitted by the exogenous fluorescent agent agent ).

[0264] Without being limited to any particular theory, the fluorescence measurements obtained by the system 200 for determining renal function include emission wavelength photons detected by the second (filtered) light detector 224. These emission wavelength photons are emitted by an exogenous fluorescent agent introduced into the patient's tissue in response to illumination with excitation wavelength photons. The emission wavelength photons travel from the fluorescence source (i.e., the exogenous fluorescent agent) through the third region 210 of the patient's skin to the second (filtered) light detector 224. However, the emission wavelength light detected by the second (filtered) light detector 224 may also include autofluorescence emitted by endogenous chromophores (such as keratin and collagen) within the patient's tissue, as well as leakage of excitation wavelength light through the optical filter 244 of the second light detector 224. The excitation wavelength photons that induce fluorescence of the exogenous fluorescent agent are generated by the first light source 218 and are directed into the first region 206 of the patient's skin. If the optical properties (scattering and / or absorption) of the patient's skin change over the time interval over which the detector data used to determine renal function is collected (i.e., from a few hours to approximately 24 hours or longer), the accuracy of the fluorescence measurement may be affected, as discussed above.

[0265] During each measurement cycle in one aspect, system 200 can direct light into a first region 206 of the patient's skin with an alternating series of emission wavelength light pulses and excitation wavelength light pulses, and can detect all of the light emitted from a second region of the patient's skin using a first (unfiltered) light detector 222, and detect a portion of the light emitted from a third region 210 of the patient's skin using a second (filtered) light detector 224. The light intensity detected by each combination of excitation and emission wavelength illumination of first region 206 and detection by unfiltered / filtered light detectors 222 / 224 contains information not only about the concentration of exogenous fluorescent agents in the patient's tissue, but also about the optical properties of the patient's skin.

[0266] Table 2: Photodetector measurements corrected for temperature and power fluctuations

[0267]

[0268] The primary measure of fluorescence is Flr meas , the fluorescence light intensity measured at the filtered detector.

[0269] Diffuse reflectance measurement Flr meas represents the propagation of photons to the unfiltered device and consists primarily of excitation photons.

[0270] DR em and DR em,filtered represents the propagation of emitted photons only.

[0271] Referring to Table 2, in various aspects, the light intensity measured by the second (filtered) light detector 224 during illumination with the excitation wavelength light, before any correction for tissue optical properties obtained by the exogenous fluorescent agent (Flr meas ) is the raw intensity of the light emitted. After baseline subtraction correction as previously described herein, it is assumed that the light contained in Flr meas The emission wavelengths in the spectrum are derived primarily from exogenous fluorophores, with only a small contribution from autofluorescence of endogenous fluorophores, and are therefore referred to as Flr agent In one aspect, if it is assumed that the optical properties of the patient's skin do not change, all autofluorescence contributions will be subtracted during the baseline correction described above.

[0272] However, if the optical properties of the patient's skin change during data acquisition, the patient's skin may exhibit slightly more or less autofluorescence at the emission wavelength, thereby introducing uncertainty into the accuracy of the previously performed background subtraction correction. Furthermore, the changing optical properties of the skin may further alter the intensity of light reaching the exogenous fluorescer at the excitation wavelength, thereby altering the amount of energy absorbed by the exogenous fluorescer and the intensity of the induced fluorescence emitted from the exogenous fluorescer in response to illumination with the excitation wavelength light. In various aspects, the remaining three optical measurements are capable of monitoring the optical properties of the patient's skin and providing data that can be used to adjust for any changes in the optical properties of the patient's skin, including the effects of autofluorescence and light leakage at the excitation wavelength.

[0273] Referring again to Table 2, the light intensity measured by the first (unfiltered reference) light detector 222 during illumination by the excitation wavelength light captures a measure of the diffuse reflectance (DR) of the excitation wavelength light propagating through the patient's skin. ex ). Although the first light detector 222 is configured to detect both the excitation wavelength light and the emission wavelength light, the intensity of the excitation wavelength light is several orders of magnitude higher than the intensity of the emission wavelength light due to the low efficiency of light generation via fluorescence. In various aspects, assuming that the emission wavelength light is in the DR ex In other aspects, estimate and subtract the emission wavelength light in DR ex Without being limited to any particular theory, because the intensity of the excitation wavelength light directed into the patient's skin is assumed to be relatively constant with negligible losses due to absorption by the exogenous fluorescer, and power correction is performed as previously described herein, DR ex Used as a baseline measurement to assess changes in the optical properties of a patient's skin relative to the excitation wavelength light.

[0274] The light intensity measured by the first (unfiltered reference) light detector 222 during illumination by the emission wavelength light captures a measure of the diffuse reflection of the emission wavelength light propagating through the patient's skin (DR em Without being limited to any particular theory, because the exogenous fluorescer was not induced to emit emission wavelength light during this phase of the data acquisition cycle due to the absence of excitation wavelength illumination, and because the intensity of the emission wavelength light directed into the patient's skin was relatively constant and power corrected as previously described herein, the DR ex Used as a baseline measurement to assess changes in the optical properties of a patient's skin relative to the emitted wavelength of light.

[0275] The light intensity measured by the second (filtered) light detector 224 during illumination by the emission wavelength light obtains a second measurement of the diffuse reflectance (DR) of the emission wavelength light propagating through the patient's skin. em,filtered In one aspect, DR em,filteredCarry out the above DR em The same assumptions apply. In addition, DR em,filtered Provides a means to assess the heterogeneity of tissue optical properties. em,filtered The light is measured by the second light detector 224, which is configured to detect light emitted from the patient's skin at the third region 210 (see Figure 2 ), so in DR em,filtered The intensity of the light measured in DR has traveled along a light path through the patient's skin that is different from that in em Without being limited to any particular theory, the distances between the first detector aperture 1004 and the second optical aperture 2006 (through which light is transmitted to the first and second optical detectors 222 / 224) are designed to be equidistant from the light transmission aperture 1002, respectively (see Figure 10 ), so assuming DR em,filtered and DR em Any difference between the two is due to heterogeneity in the optical properties of the light passing through the skin via the two different paths.

[0276] Excitation wavelength leak correction

[0277] In one aspect, is used as a basis for estimating the leakage through of the excitation wavelength light into the second (filtered reference) photodetector 224, which is used as part of the method of eliminating the effects of variations in the background signal described herein. Without being limited to any particular theory, it is assumed that the amount of leakage through of the excitation wavelength light into the second (filtered reference) photodetector 224 is related to the DR ex The signal is proportional to the optical properties of the patient's skin and is affected only by factors related to the device and not by factors related to the optical properties of the patient's skin. ex The ratio is assumed to be constant.

[0278] In one aspect, the excitation wavelength light leakage (ExLT) included in the raw fluorescence signal (Flr) is assumed to be the signal according to equation (21) The constant ratio C ExLT :

[0279]

[0280] Among them C ExLT is the sensor head specific calibration factor.

[0281] In one aspect, C is obtained by calculating the ratio (Det1 / Det2) between the excitation light detected by the first and second light detectors 222 / 224 on the non-fluorescent optical phantom according to equation (22). ExLT :

[0282]

[0283] On the other hand, due to tissue inhomogeneities, the excitation light reaching the filtered detectors is assumed to be different from the light reaching the unfiltered detectors. In this regard, the ratio of the emission wavelength light at each detector is used to correct for this inhomogeneity.

[0284] In all cases, the calibration factor C ExLT may be specific to a single sensor head 204 or may be based on various factors including, but not limited to, manufacturing tolerances. ExLT can be applied to all sensor heads 204 of the system 200. In one aspect, if the system 200 is used to obtain C ExLT , Flr meas and A non-fluorescent homogeneous phantom in the environment of the system 200 as described above. It should be noted that equation (22) assumes that the tissue monitored by the system 200 is homogeneous.

[0285] In one aspect, the excitation wavelength light leakage (ExLT) determined by equation (21) can be obtained from the raw fluorescence signal (Flr meas ) to obtain the corrected fluorescence signal F as described in equation (23) photons :

[0286] Flr photons =Flr meas -ExLT Equation (23)

[0287] Figure 17A In one aspect, the raw fluorescence signal (Flr) is obtained by the system 200 before and after injection of an exogenous fluorescent agent. meas , blue line) and the corresponding excitation wavelength light leakage (ExLT, red line) determined using equation (23). Figure 17A As shown, the ExLT signal changes during the data acquisition process. Figure 17B The original fluorescence signal (Flr meas , blue line) and the fluorescence signal (Flr) after eliminating the leakage of the excitation wavelength as described above in equation (23) photons , green line) for comparison.

[0288] In one aspect, the raw fluorescence signal Flr is first corrected meas, to eliminate the effect of light leakage at the excitation wavelength using equation (23). In this regard, as described below, the corrected fluorescence signal Flr is used photons It serves as a basis for subsequent correction to eliminate the influence of autofluorescence.

[0289] Fluorescence leakage correction

[0290] Without being limited to any particular theory, the light detected by the unfiltered light detector during illumination with light at the excitation wavelength is a mixture of diffuse reflection of the excitation wavelength and light from the reagent fluorescence. In one aspect, it is assumed that the diffuse reflection is sufficiently stronger than the fluorescence so that the contribution of the fluorescence to the unfiltered detector measurement is negligible.

[0291] On the other hand, the contribution of fluorescence to the unfiltered detector measurement is non-negligible. As a non-limiting example, Figure 27 It is shown that without the application of exogenous fluorescent agents for more than a day and Flr meas However, if Figure 28 As shown, The signal occasionally showed fluorescence leakage, as evidenced by an associated signal increase after the reagent entered the patient's bloodstream.

[0292] On the one hand, according to equation (24) from the original signal Eliminate the part of the diffuse excitation signal caused only by the excitation photons:

[0293]

[0294] In each case, the fluorescence leakage Flr is measured leakt hroug h Quantity and Signal The relationship between The coefficient C was determined empirically based on the tissue heterogeneity expressed by FlrLT In one aspect, measurements may be obtained from multiple subjects. As a non-limiting example, Figure 29 is the Flr determined by general experience leakt hroug h and from a database of 33 patients In this regard, the empirically derived relationship was examined across multiple patient datasets and found to be consistent. The correction factor ExLT was set to include the relationship between tissue heterogeneity and fluorescence leakage flux and is defined as follows:

[0295]

[0296] In one aspect, equation (25) includes p1=0.6138 and p2=0.01095, as determined by Figure 29Determined by a biquadratically weighted linear fit of the relationships shown.

[0297] In another aspect, C is determined by obtaining measurements of an optical model provided with increasing fluorescence concentrations. FlrLT , where the only variation in signal is due to the concentration of the exogenous fluorescer.

[0298] Separation of fluorescence and diffuse reflectance at excitation wavelengths

[0299] In all aspects, DR ex The number of photons on a filtered or unfiltered detector due to Flr depends on the directionality of the light and the gain of each detector at the detection wavelength as follows:

[0300] DRex meas =A1*DRex photons +B1*Flr photons Equation (26)

[0301] Flr meas =A2*DRex photons +B2*Flr photons Equation (27)

[0302] Where coefficients A1, A2, B1 and B2 include directivity and gain factors. As a non-limiting example, A1 can be set in the form of equation (28):

[0303] A1=d450 SPM1 *G SPM1@450 Equation (28)

[0304] Among them d450 SPM1 and G SPM1@450 are the directivity and gain factor of the detector SPM1 at an illumination wavelength of 450 nm.

[0305] In one aspect, the photon signal can be isolated as shown in Equation (29) and Equation (30):

[0306]

[0307]

[0308] In various aspects, a constant term (such as ) is not needed because the renal function monitor as disclosed herein measures the rate of change of intrinsic fluorescence (IF) expressed by equation (31):

[0309]

[0310] In one aspect, as described above, the item (or C ExLT )and (or C FlrLT ) was determined experimentally to separate Flr photons and DRex photons .

[0311] Autofluorescence correction

[0312] In various aspects, the method of correcting measured fluorescence to eliminate time-varying effects of background can also include eliminating autofluorescence effects other than excitation wavelength leakage effects. As used herein, autofluorescence refers to the emission wavelength light produced by endogenous chromophores (such as keratin and collagen) in response to irradiation with excitation wavelength light. In various aspects, autofluorescence can change during the process of obtaining fluorescence measurements using the systems and methods described herein. Without being limited to any particular theory, changes in the optical properties of the patient's skin, such as changes in the concentration of chromophores (such as hemoglobin and / or melanin), can cause changes in the level of autofluorescence.

[0313] Figure 18 is the sum of the raw fluorescence obtained during the background interval (Flr meas , blue line), defined herein as the interval before the exogenous fluorescent agent is injected into the patient 202 when it is assumed that the patient tissue does not contain the exogenous fluorescent agent. Figure 18 Also shown is the use of equation (5) from Flr meas The signal generated by eliminating the effect of leakage of the excitation wavelength light (ExLT) in the background interval, as described above. The remaining fluorescence signal detected during the background interval, such as Figure 18 As shown in the green line, various aspects can be assumed to be attributed to autofluorescence.

[0314] In one aspect, intrinsic autofluorescence (IF auto ), defined herein as the measured fluorescence at the emission wavelength that is attributed solely to the emission of endogenous chromophores (such as keratin and collagen), can be calculated according to equation (32) as the corrected fluorescence signal Flr obtained during the background interval photons (See Equation 23).

[0315] IF Auto =median(Flr photons (1:endBackground)) Equation (32)

[0316] where endBackground is the index of the data collection in the dataset that corresponds to the end of the background interval just before the injection of the exogenous fluorescent agent.

[0317] On the one hand, it can be assumed that the autofluorescence is relatively stable throughout the data acquisition process, including the interval after the injection of the exogenous fluorescent agent. In this regard, as shown in Equation (33), the fluorescence signal Flr can be obtained from the corrected fluorescence signal photons Subtract the IF obtained in equation (32) from auto The value of is used to eliminate the influence of autofluorescence:

[0318] IF agent =Flr photons -IF Auto Equation (33)

[0319] Where IF agent Intrinsic fluorescence refers to the emission wavelength of light emitted by exogenous fluorescent agents.

[0320] Figure 19A is a graph summarizing the various measurements obtained during the background interval: raw fluorescence (Flr meas ), (red line), DR em (orange line) and DR em,filtered (purple line). In addition, the intrinsic autofluorescence (IF auto , green line) is also Figure 19A As shown in Figure 19A During the background interval shown, the values of all quantities are relatively stable.

[0321] Figure 19B It is a summary Figure 19A Graph of diffuse reflectance measurement shown: (red line), DR em (orange line) and DR em,filtered (Purple line). During the process of obtaining fluorescence measurements after injection of exogenous fluorescent agents (i.e., Figure 18 As shown, after a time of approximately 9:07, the diffuse reflectance measurement decreases significantly, indicating that the optical properties of the patient's skin that affect the measurement signal from autofluorescence may also change during this time period.

[0322] In another aspect, diffuse reflectance measurements can be used to project the underlying autofluorescence signal over the entire measurement period, thereby accounting for changes in the optical properties of the patient's skin during the entire process of data measurement. In one aspect, diffuse reflectance measurements can be used to scale the corrected fluorescence signal Flr photons In order to take into account the variations in the optical properties of the patient's skin, an intrinsic fluorescence is generated. In this regard, in order to correct the fluorescence measurements obtained after the injection of an exogenous fluorescent agent, the combined intrinsic fluorescence IF can be obtained from equation (33): AgentAndAuto The intrinsic autofluorescence (IF) calculated by equation (32) is subtracted from auto), as shown in equation (34):

[0323] IF agent =IF AgentAndAuto -IF Auto Equation (34)

[0324] In one aspect, the background correction subunit 1306 can enable Figure 20 The background correction method 2000 is summarized in the block diagram of FIG. The method 2000 may include performing a correction at 2002 to eliminate the effects of leakage of emission wavelength light into the second (filtered reference) light detector 224, as described in equations (29), (30), and (31). The method 2000 may also include estimating the autofluorescence (IF) at 2004 from analysis of measurements obtained during the background interval, as described in equation (32) above. auto ) level. The method 2000 may also include performing a correction at 2006 to remove the effect of autofluorescence from the fluorescence measurement, as described in equation (33) above. In fact, the autofluorescence signal IF auto is forward-projected into the subsequent fluorescence measurement and is removed at 2006. As described below, the intrinsic fluorescence IF generated by the background effect is removed from the original fluorescence measurement. agent , which may be converted by the RDTC calculation subunit 1310 into parameters including but not limited to glomerular filtration rate (GFR) and / or renal decay time constant (RDTC) representing renal function.

[0325] e) Fault detection subunit

[0326] Refer again Figure 13 The processing unit 236 of the controller 212 may also include a fault detection subunit 1312 configured to monitor the functionality of the light sources 218 / 220 and the light detectors 222 / 224 and notify the user of any irregularities in any detected faults within the system 200 via the display unit 216. In various aspects, the fault detection subunit 1312 may detect any irregularities in any detected faults by examining the signal levels received from the light sources 218 / 220 and the light detectors 222 / 224 and the associated sensor heads 204 (see FIG. Figure 2) and an additional light detector 226 to enable basic identification of fault and notification conditions. In various aspects, the signal amplitude (see equation (1)) and the average signal can be used to determine the peak and nadir levels of the modulation of the LED light source 218 / 220. The nadir of the signal, defined herein as the average signal minus half the peak-to-peak signal, can be used to monitor the ambient light level in one aspect. Without being limited to any particular theory, additional contributions to the nadir level of the modulated signal, such as amplifier DC offset, can be neglected to be small and constant relative to the contribution of ambient light leakage. In one aspect, if the ambient light level detected at low detector amplifier gain records more than about one-quarter of the range of the high speed ADC1102, an ambient light notification is issued to the user via the display unit 216.

[0327] In various other aspects, saturation of the light detectors 222 / 224 can also be monitored by the fault detection subunit 1312. In these other aspects, saturation can be monitored by calculating the peak value of the signal, defined herein as the average signal value plus half the peak-to-peak signal. If the peak value of the signal falls within 5% of the saturation range of the ADC, the fault detection subunit 1312 can issue a saturation notification to the user via the display unit 216. If the fault detection subunit 1312 detects a saturation event, it can then check the ambient light level to determine whether the saturation event is associated with ambient light saturation, defined herein as a saturation event occurring simultaneously with the ambient light notification described above. If an ambient light saturation event is detected, the fault detection subunit 1312 issues an ambient light saturation notification to the user via the display unit 216, and data acquisition by the acquisition unit 234 continues in this notification state to allow the user to address the situation. If a saturation event not associated with excess ambient light is detected, the fault detection subunit 1312 can signal the light detector control unit 232 to perform an adjustment of the detector gain and / or can signal the light source control unit 230 to perform an adjustment of the LED current source 1126 to adjust the LED intensity. In various aspects, the fault detection subunit 1312 issues a notification to the user via the display unit to report the ambient light saturation event or the saturation event not associated with excess ambient light. In some aspects, if a saturation event is detected, but the user has disabled automatic gain control when the system 200 is configured in engineering mode as described above, the user is also notified via the display unit.

[0328] e) Post-agent administration selection unit

[0329] Refer again Figure 13 The processing unit 236 may further include a post-reagent application subunit 1308 configured to automatically identify a portion of the measurement data set corresponding to a post-reagent application region, as described below.

[0330] Figure 21 is a graph of fluorescence measurements obtained from a patient over a period of about 10 hours following injection of an exogenous fluorescent agent (such as MB-102), following a pre-injection period 2102 of about 3 hours. Figure 21 , the pre-injection / baseline period 2102 is characterized by relatively low and stable fluorescence levels, which may be due to the absence of exogenous fluorescent agent in the patient's blood. After the injection 2103 of the exogenous fluorescent agent, the fluorescence measurement shows a sharp increase 2106 to a peak concentration 2108, followed by a relatively smooth exponential decrease back to background fluorescence levels as the kidneys eliminate the exogenous fluorescent agent from the patient's blood. Without being limited to any particular theory, it is believed that after the amount of time for the exponential concentration decrease has passed, the injected exogenous fluorescent agent may be well mixed.

[0331] Refer again Figure 21 After an exogenous fluorescent agent, such as MB-102, is injected into a patient's bloodstream, the exogenous fluorescent agent undergoes an equilibrium period of diffusion from the bloodstream into the patient's remaining extracellular tissues. After the agent is injected 2103, the temporal distribution of the fluorescence signal IF can be characterized as a bi-exponential signal distribution described by equation (35):

[0332]

[0333] where C0 is the baseline signal which is typically removed by baseline subtraction as described above.

[0334] Refer again Figure 21 Once the diffusion of the exogenous fluorescent agent into the patient's extracellular tissue reaches a quasi-steady-state condition, post-equilibrium 2110 is reached and the fluorescence signal can be characterized as a linear decay. Without being limited to any particular theory, assume that the post-equilibrium region 2110 of the measurement data set is characterized as the time region of the IF data set, when logarithmically transformed, the region of the IF time distribution is well described by a linear equation. In one aspect, the post-equilibrium region is well described by equation (36):

[0335] IF post-equilibration =C0+C1e -t / τ Equation (36)

[0336] In one aspect, the post-reagent administration selection subunit 1308 can automatically identify the post-equilibration period 2110 by performing a single exponential curve fit on different portions of the IF dataset and analyzing the associated curve fit error for each of the different portions. In various aspects, the post-reagent administration selection subunit 1308 can select the earliest occurring portion of the IF dataset where the curve fit error associated with the single exponential curve fit is below a threshold as the initial post-equilibration portion of the IF dataset suitable for data correction and analysis as described above. Any analysis method suitable for comparing the curve fit errors associated with single exponential curve fits of different portions of the IF dataset can be used in the post-reagent administration selection subunit 1308, including but not limited to linear curve fit portions of the IF dataset falling within overlapping or non-overlapping data windows and comparing the curve fit errors of the corresponding data windows. In one aspect, the post-reagent administration selection subunit 1308 can generate at least one signal configured to signal the RDTC calculation subunit 1310 for a time range within the IF dataset corresponding to the post-equilibration period 2110, thereby enabling selection of an appropriate portion of the IF dataset for correction and analysis, as disclosed herein.

[0337] In another aspect, a linear fit and a 2-exponential fit to the IF data can be compared. In this other aspect, equilibrium can be identified as complete once the fit errors are equivalent (correcting for the extra degrees of freedom in the 2-exponential fit).

[0338] f) RDTC calculation subunit

[0339] In various aspects, the system 200 is configured to convert various measurements from the photodetectors 222 / 224 and associated light sources 218 / 220, as well as other thermal and optical sensors, in response to illumination with light at the excitation wavelength, into a corrected intrinsic fluorescence (IF) signal corresponding to the detected fluorescence, which is attributable solely to fluorescence emitted by the exogenous fluorescent agent at the emission wavelength. In various aspects, the exponential decrease in the IF signal during the post-agent administration portion of the IF dataset can be analyzed to monitor and quantify renal function.

[0340] In one aspect, the exponential decline of IF signal during the rear reagent administration portion of IF data set can be converted into glomerular filtration rate (GFR), which is configured to quantify renal function. In yet another aspect, the exponential decline of IF signal during the rear equilibrium portion of IF data set can be converted into kidney decay time constant (RDTC), which is also configured to quantify renal function. In yet another aspect, the exponential decline of IF signal during the rear equilibrium portion of IF data set can be converted into kidney decay rate, which is also configured to quantify renal function.

[0341] Refer again Figure 13 , the processing unit 236 may further include an RDTC calculation subunit 1310, which is configured to automatically convert the IF signal into a renal decay time constant (RDTC). As used herein, the renal decay time constant (RDTC) is defined as the time constant associated with the post-equilibrium monoexponential decay described in equation (36) above. In one aspect, after accurate baseline subtraction is performed by the baseline subtraction subunit 1304, the renal decay time constant τ can be calculated by performing linear regression on the logarithmic transformed IF signal data (log(IF)), as described in equation (37):

[0342]

[0343] In various aspects, the RDTC calculation subunit 1310 can generate a signal configured to generate a display of the calculated RDTC using the display unit 216. The display of the calculated RDTC can be provided to the display unit 216 in any suitable format, including, but not limited to, a graph of the RDTC as a function of time, a single discrete RDTC value, a table of RDTC values as a function of time, a color-coded display, or other graphical representation configured to determine whether the calculated RDTC can be classified as normal / healthy, abnormal, high, low, or any other suitable classification. In various other aspects, any of the aforementioned graphical formats can be continuously or non-continuously updated as additional data is acquired and analyzed. In one aspect, the RDTC calculation subunit 1310 can calculate the RDTC as described above within non-overlapping and / or overlapping windows within the IF data set.

[0344] On the other hand, the RDTC calculation subunit 1310 can convert the RDTC into a glomerular filtration rate (GFR) using known methods. In this regard, the RDTC can be inverted and multiplied by the slope to produce cGFR, a prediction of GFR that can be corrected for body size (e.g., body surface area or volume of distribution).

[0345] v) Memory

[0346] Refer again Figure 2Controller 212 of system 200 may also include memory 242, which is configured to facilitate data storage within system 200. In some embodiments, memory 242 includes multiple storage components, such as, but not limited to, a hard drive, flash memory, random access memory, and a magnetic or optical disk. Alternatively or additionally, memory 242 may include a server, such as a remote storage device, in communication with controller 212. Memory 242 stores at least one computer program that, when received by at least one processor, causes the at least one processor to perform any of the functions of controller 212 described above. In one embodiment, memory 242 may be or include a computer-readable medium, such as a floppy disk, hard disk, optical disk, or magnetic tape device, flash memory or other similar solid-state memory device, or an array of devices, including devices in a storage area network or other configuration. A computer program product may be tangibly embodied in an information carrier. A computer program product may also include instructions that, when executed, perform one or more functions, such as those described herein. The information carrier may be a non-transitory computer or machine-readable medium, such as memory 242 or memory on processor 238.

[0347] In various aspects, the system 200 can log raw measurements and processed data into a series of files. Each file can include a header file containing information about the operator, tool, and session. Each experimental session logs a set of files into separate folders for each sensor head used in that session. The raw data files can contain in-phase, quadrature, and average measurements from the detector and monitor during the active period of the excitation and emission wavelength LEDs, as well as the gain settings of the LEDs and detectors at the time of data acquisition.

[0348] In various other aspects, the processed data file can contain fluorescence and diffuse reflectance measurements after amplitude calculation and correction of the monitor readings, as well as the gain settings of the LED and detector. The intrinsic fluorescence data file can contain intrinsic fluorescence measurements resulting from diffuse reflectance correction of the raw fluorescence signal. The GFR file can contain calculated GFR as a function of time, categorized to indicate whether post-equilibration has occurred, and confidence limits. The telemetry file can contain temperature and voltage measurements. The event log file can contain user and automatically generated event records.

[0349] vi) GUI unit

[0350] Refer again Figure 2In various aspects, the controller 212 can include a GUI unit 240 configured to receive a plurality of signals encoding various measured and transformed data from other units of the system. Furthermore, the GUI unit can be configured to generate signals configured to operate the display unit 216 to display data, frames, tables, and / or any other communication of information between a user and the system 200.

[0351] vii) Processor

[0352] Refer again Figure 2 , the controller 212 may also include a processor 238. The processor 238 may include any type of conventional processor, microprocessor, or processing logic that interprets and executes instructions. The processor 238 may be configured to process instructions for execution within the controller 212, including instructions stored in the memory 242 for displaying graphical information of a GUI on an external input / output device (such as a display unit 216 coupled to a high-speed interface). In other implementations, multiple processors and / or multiple buses as well as multiple memories and multiple types of memories may be used as appropriate. In addition, multiple controllers 212 may be connected to each device, each of which provides a portion of the necessary operations to enable the functionality of the system 200. In some embodiments, the processor 238 may include an acquisition unit 234, a light detector control unit 232, a light source control unit 230, and / or a processing unit 236.

[0353] As used herein, a processor such as processor 238 may include any programmable system, including systems using microcontrollers, reduced instruction set circuits (RISC), application specific integrated circuits (ASICs), logic circuits, and any other circuits or processors capable of performing the functions described herein. The above examples are merely examples and are therefore not intended to limit the definition and / or meaning of the term "processor" in any way.

[0354] As described herein, computing devices and computer systems include processors and memory. However, any processor in a computing device referred to herein may also refer to one or more processors, where the processor may be in a single computing device or in multiple computing devices operating in parallel. Additionally, any memory in a computing device referred to herein may also refer to one or more memories, where the memory may be in a single computing device or in multiple computing devices operating in parallel.

[0355] C.Operation unit

[0356] The operation unit 214 may be configured to enable a user to interface with the controller 212 (e.g., visually, audio, touch, button presses, stylus taps, etc.) to control the operation of the system 200. In some embodiments, the operation unit 214 may be further coupled to each sensor head 204 to control the operation of each sensor head 204.

[0357] D.Display unit

[0358] Refer again Figure 2 The system 200 may also include a display unit 216 configured to enable a user to view data and control information for the system 200. The display unit 216 may also be coupled to other components of the system 200, such as the sensor head 204. The display unit 216 may include a visual display, such as a cathode ray tube (CRT) display, a liquid crystal display (LCD), a light emitting diode (LED) display, or an "electronic ink" display. In some embodiments, the display unit 216 may be configured to present a graphical user interface (e.g., a web browser and / or client application) to the user. The graphical user interface may include, for example, a display of the GFR values generated by the system 200, as described above, as well as operational data for the system 200.

[0359] Exogenous markers

[0360] Without being bound by any particular theory, molecules with high hydrophilicity and small (creatinine, molecular weight = 113) to medium size (inulin, molecular weight approximately 5500) are known to be rapidly cleared from the systemic circulation by glomerular filtration. In addition to these properties, an ideal GFR agent would be neither reabsorbed nor secreted by the renal tubules, would have negligible binding to plasma proteins, and would have very low toxicity. To design an optical probe that meets all of these requirements, a balance is struck between photophysical properties and the molecular size and hydrophilicity of the fluorophore. For example, although hydrophobic cyanine and indocyanine dyes absorb and emit optimally in the near-infrared (NIR) biological window (700 to 900 nm), they are not hydrophilic enough to be used as pure GFR agents. Smaller dye molecules may be more easily converted to the extremely hydrophilic species required for renal clearance, but due to the limited π systems formed by these lower molecular weight compounds, they are generally able to be excited and emit single photons in the ultraviolet (UV).

[0361] To address the pharmacokinetic challenges associated with enhanced photophysical properties, simple derivatives of 2,5-diaminopyrazine-3,6-dicarboxylic acid have been used as very low molecular weight fluorescent scaffold systems with bright emission in the yellow to red region of the electromagnetic spectrum. To simultaneously optimize both pharmacokinetic and photophysical properties for GFR, SAR studies have been conducted using amide-linked variants of these derivatives. A variety of hydrophilic functionalities can be employed to achieve rapid renal clearance of this class of pyrazine fluorophores, including carbohydrates, alcohols, amino acids, and various PEG-based strategies. PEG surrogates can be used to increase hydrophilicity and solubility, reduce toxicity, and modulate the aggregation of the resulting pyrazine derivatives. A range of medium-sized PEG-pyrazine derivatives, varying in molecular weight and structure (and therefore hydrodynamic volume), are also suitable for use as intrinsic fluorescers.

[0362] In one aspect, the exogenous fluorescent agent is MB-102.

[0363] Example

[0364] The following examples illustrate various aspects of the disclosed systems and methods.

[0365] Example 1: Sensor head with flared housing

[0366] Figure 23 FIG2 is a perspective view of another aspect of sensor head 204a. In this aspect, sensor head 204a includes a housing 600a formed by an upper housing 602a and a flared lower housing 604a. The surface area of lower housing 604a is enlarged to form an enlarged bottom surface 608a. Housing 600a also includes a cable opening 806a formed through upper housing 602a.

[0367] Figure 24 6 is a bottom view of the sensor head 204a showing the bottom surface 608a of the housing 600a. The bottom surface 608a may include an aperture plate 702a including one or more apertures 704a configured to transmit light between the patient's skin and the light source and light detector contained within the housing 600. Figure 24 As shown, aperture 704a includes a light delivery aperture 1002a configured to deliver illumination generated by first and second light sources 218 / 220 to tissue of patient 202, and first and second detector apertures 1004 / 1006 configured to receive light from tissue of patient 202. In one aspect, bottom surface 608a enables positioning aperture 704a below a relatively large area that is shielded from ambient light conditions by bottom surface 608a. This reduction in scattered ambient light entering first and second detector apertures 1004 / 1006 reduces noise introduced into light intensity measurements obtained by first and second light detectors 222 / 224.

[0368] In various aspects, the bottom surface 608a of the housing 600a can be attached to the patient's skin using a biocompatible and transparent adhesive material 610a, including but not limited to a transparent double-sided medical grade adhesive, such as Figure 24 A transparent adhesive material 610a can be positioned on the bottom surface 608a such that the adhesive material 610a covers the hole 704a.

[0369] Figure 25 is an isometric view of the sensor head 204a with the upper housing 602a and various electrical components removed to expose the inner housing 2502. Figure 26 yes Figure 25 An exploded view of the inner housing 2502 and associated electrical components is shown. Figure 25 and Figure 26 , inner housing 2502 is contained within housing 600a and mounted to lower housing 608a. Inner housing 2502 includes a sensor mount 912 having a first detection well 908, a second detection well 910, and a light source well 902 formed therethrough. A first light detector 222 is mounted within first detection well 908, and a second light detector 224 is mounted within second detection well 910. First and second light sources 218 / 220 are mounted within light source well 902. In one aspect, first detection well 908, second detection well 910, and light source well 902 of sensor mount 912 are optically isolated from one another to ensure that light from light sources 218 / 220 does not reach light detectors 222 / 224 without being coupled through the skin of patient 202. As described in detail above, the separation between the two detection wells 908 / 910 ensures that the detected fluorescence signal from the exogenous fluorescer is distinguishable from unfiltered excitation light.

[0370] Reference Figure 26 , inner housing 2502 includes a first detection hole 2602, a second detection hole 2604, and a light source hole 2606. Sensor mount 912 is coupled to inner housing 2502 such that first detection hole 2602, second detection hole 2604, and light source hole 2606 are aligned with first detection well 908, second detection well 910, and light source well 902 of sensor mount 912, respectively.

[0371] In one aspect, optically transparent windows 2610, 2612, and 2614 are coupled within the first detection aperture 2602, the second detection aperture 2604, and the light source aperture 2606, respectively, to seal the apertures while also providing an optically transparent conduit between the tissue and the interior of the sensor head 204a. Furthermore, diffusers 2616, 2618, and 2620 are coupled to the optically transparent windows 2610, 2612, and 2614, respectively. Diffusers 2616, 2618, and 2620 are configured to spatially homogenize the light transmitted to the tissue by the light sources 218 / 220 and to spatially homogenize the light detected by the light detectors 222 / 224. In one aspect, an absorption filter 244 is coupled to the diffuser 2616. In one aspect, the absorption filter 244 is coupled to the diffuser 2616 using an optically transparent adhesive.

[0372] In view of the foregoing, it can be seen that the multiple advantages of the present disclosure are achieved and other favorable results are obtained. As various changes can be made to the above-described methods and systems without departing from the scope of the present disclosure, it is intended that all matter contained in the above description and shown in the accompanying drawings shall be interpreted as illustrative and not in a restrictive sense.

[0373] When introducing elements of the present disclosure or versions, embodiments, or aspects thereof, the articles "a," "an," "the," and "said" are intended to mean that there are one or more of the elements. The terms "comprising," "including," and "having" are intended to be inclusive and mean that there may be additional elements other than the listed elements.

Claims

1. A method for monitoring time-varying fluorescence emitted from a fluorescent agent within a diffusely reflecting medium having time-varying optical properties, the method comprising: A monitoring system is provided, wherein a measurement data set is provided, comprising a plurality of measurement data entries, each measurement data entry comprising at least two measurement values obtained from a patient at a data acquisition time before and after administration of an exogenous fluorescent agent, wherein the at least two measurement values comprise a fluorescence emission signal Flr detected by a filtered light detector at a third region adjacent to the diffusely reflecting medium during illumination of the diffusely reflecting medium by light of an excitation wavelength from a first region, and at least one diffuse reflectance signal DR, wherein the at least one diffuse reflectance signal DR is selected from: DR ex a signal detected by an unfiltered light detector at a second region adjacent to the diffusely reflective medium during illumination of the diffusely reflective medium by the excitation wavelength light from the first region adjacent to the diffusely reflective medium; DR em a signal detected by the unfiltered light detector at the second region during the period when the emission wavelength light from the first region illuminates the diffusely reflecting medium; and DR em,filtered a signal detected by the filtered light detector at the third region during the period when the emission wavelength light from the first region illuminates the diffuse reflective medium; as well as identifying, by the monitoring system, a post-agent administration portion of the measurement data set; The monitoring system converts the fluorescence emission signal Flr of each measurement data entry within the post-reagent application portion of the measurement data set into a corrected fluorescence signal, wherein the conversion includes eliminating the effect of leakage of excitation wavelength light into the fluorescence emission signal Flr and eliminating at least one of the autofluorescence contribution from the fluorescence emission signal Flr.

2. The method according to claim 1, wherein The at least one diffuse reflection signal DR is DR ex signal, and wherein eliminating the effect of leakage through of the excitation wavelength light into the fluorescence emission signal Flr comprises converting the DR ex The signal is converted into an excitation wavelength light leakage signal ExLT indicating the level of excitation wavelength light leakage: ExLT=C ExLT *DR ex Equation (21) Among them C ExLT is the calibration factor.

3. The method according to claim 2, wherein: Eliminating the effect of leakage of the excitation wavelength light into the fluorescence emission signal Flr also includes converting the fluorescence emission signal Flr into a corrected fluorescence signal representing only the detected emission wavelength fluorescence using equation (23) Flr photons ： Flr photons =Flr-ExLT equation (23).

4. The method according to claim 2, wherein: Eliminating the autofluorescence contribution comprises determining IF by determining a value of the fluorescence emission signal Flr over a portion of the measurement data set obtained before administering the exogenous fluorescent agent. auto , the IF auto represents intrinsic autofluorescence emitted by endogenous chromophores within the diffusely reflecting medium, said value being selected from the group consisting of a mean and a median.

5. The method according to claim 4, wherein Eliminating the autofluorescence contribution further comprises subtracting IF from the fluorescence emission signal Flr. auto .

6. The method according to claim 2, wherein: C ExLT Obtained by the following steps: A plurality of measurements are obtained from the solid model, the plurality of measurements comprising: The fluorescence emission signal Flr represents the emission wavelength fluorescence measured using the filtered light detector; The excitation wavelength light signal DR ex , measured using an unfiltered photodetector; and Calculate C according to equation (22) ExLT :

7. A method of determining renal function in a patient, the method comprising: A monitoring system is provided, wherein a measurement data set is provided, comprising a plurality of measurement entries, each measurement data entry comprising at least two measurement values obtained from tissue of a patient at corresponding data acquisition times before and after administration of an exogenous fluorescent agent, wherein the at least two measurement values comprise a fluorescence emission signal Flr detected by a filtered light detector at a third region adjacent to the diffusely reflecting medium during illumination of the diffusely reflecting medium by light of an excitation wavelength from a first region, and at least one diffuse reflectance signal DR, wherein the at least one diffuse reflectance signal DR is selected from: DR ex a signal detected by an unfiltered light detector at a second region adjacent to the diffusely reflective medium during illumination of the diffusely reflective medium by the excitation wavelength light from the first region adjacent to the diffusely reflective medium; DR em a signal detected by the unfiltered light detector at the second region during the period when the emission wavelength light from the first region illuminates the diffusely reflecting medium; and DR em,filtered a signal detected by the filtered light detector at the third region during the period when the emission wavelength light from the first region illuminates the diffusely reflecting medium; as well as identifying a post-agent administration portion of the measurement data set; The fluorescence emission signal Flr of each measurement data entry within the post-reagent administration portion of the measurement data set is converted into an intrinsic fluorescence signal IF by a monitoring system. agent , the IF agent represents the corrected fluorescence intensity emitted by the exogenous fluorescent agent from within the diffuse reflecting medium; identifying, by a monitoring system, a post-balanced portion of the measurement data set; The intrinsic fluorescence signal IF from the post-equilibration portion of the measurement data set is used by a monitoring system agent The rate of change determines renal function.

8. The method according to claim 7, wherein: The conversion of the fluorescence emission signal Flr includes at least one of the following: eliminating the effect of leakage of the excitation wavelength light into the fluorescence emission signal Flr and eliminating the autofluorescence contribution from the fluorescence emission signal Flr, wherein the at least one diffuse reflection signal DR is DR ex signal, and wherein eliminating the effect of leakage of the excitation wavelength light into the fluorescence emission signal Flr comprises: using equation (21) to convert the DR ex The signal is converted into an excitation wavelength light leakage signal ExLT indicating the level of excitation wavelength light leakage: ExLT=C ExLT *DR ex Equation (21) Among them, C ExLT is the calibration factor.

9. The method according to claim 8, wherein Eliminating the effect of leakage of the excitation wavelength light into the fluorescence emission signal Flr also includes: using equation (23) to convert the fluorescence emission signal Flr meas Converted into a corrected fluorescence signal that represents only the fluorescence at the detected emission wavelength Flr photons ： Flr photons =Flr-ExLT equation (23).

10. The method according to claim 8, wherein Eliminating the autofluorescence contribution comprises determining the Flr over a portion of the measurement data set obtained before administering the exogenous fluorescent agent. photons The value of the signal determines the IF auto , the IF auto represents intrinsic autofluorescence emitted by endogenous chromophores within the diffusely reflecting medium, said value being selected from the group consisting of a mean and a median.

11. The method according to claim 10, wherein: Eliminating the autofluorescence contribution also includes photons Subtract IF auto To obtain the intrinsic fluorescence signal IF agent .

12. The method according to claim 8, wherein C ExLT Obtained by the following steps: A plurality of measurements are obtained from the solid model, the plurality of measurements comprising: a fluorescence emission signal Flr representing the emission wavelength fluorescence measured using the filtered light detector; Excitation wavelength optical signal DR ex , measured using an unfiltered photodetector; and Calculate C according to equation (22) ExLT :

13. The method according to claim 7, wherein: Convert the intrinsic fluorescence signal IF corresponding to the post-equilibrium portion of the measurement data set agent A portion of the intrinsic fluorescence signal IF is determined agent The rate of change includes: The intrinsic fluorescence signal IF is calculated according to the following equation: agent A single exponential curve fit of : IF post-equilibration =C0+C1e -t / τ Among them, IF post-equilibration represents the intrinsic fluorescence signal IF corresponding to the post-equilibrium portion of the measurement data set agent , C0 and C1 are curve fitting constants, t is time, and τ is the time constant; and Assign the time constant τ as the renal decay time constant, wherein the renal decay time constant represents the intrinsic fluorescence signal IF agent The rate of change.

14. The method according to claim 7, wherein: Convert the intrinsic fluorescence signal IF corresponding to the post-equilibrium portion of the measurement data set agent A portion of the intrinsic fluorescence signal IF agent The rate of change includes: At each corresponding data acquisition time, the intrinsic fluorescence signal IF agent Perform logarithmic transformation; The intrinsic fluorescence signal IF after logarithmic transformation is performed according to the corresponding data acquisition time agent Linear regression of to obtain the slope; Wherein, the slope represents the intrinsic fluorescence signal IF agent The rate of change.

Images