Sensor head

Abstract

This application relates to a sensor head. The sensor head includes: a housing configured to be attached to a body surface of a patient and including a bottom surface; at least two light sources configured to transmit light into a first region of the patient, the at least two light sources including a first blue LED light source and a second green LED light source, wherein the first blue LED light source transmits light at an excitation wavelength; and at least one light detector configured to detect light at an emission wavelength at a second region of the patient; wherein the housing encloses the at least two light sources and the at least one light detector.

Description

[0001] This application is a divisional application of the Chinese national phase application of the PCT application with the application date of January 30, 2018, the international application number of PCT / US2018 / 016041, and the invention title of "Non-invasive Monitoring Method Using Fluorescent Tracer with Diffuse Reflection Correction". The entry date of this Chinese national phase application into the Chinese national phase is July 26, 2019, and the application number is 201880008812.7, the entire content of which is incorporated herein by reference.

[0002] Cross-reference to related applications

[0003] This application claims the benefit of U.S. Provisional Application No. 62 / 452,025, filed on January 30, 2017, the entire content of which is incorporated herein. Technical field

[0004] The present disclosure generally relates to methods for non-invasive monitoring of fluorescent tracers within 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 within tissue in a patient's body. Background art

[0005] To minimize the risk of acute renal failure caused by various clinical, physiological, and pathological conditions, it is highly desirable to perform real-time dynamic monitoring of a patient's renal function at the bedside. This is particularly important in the case of critically ill or injured patients, as a large proportion of these patients are at risk of multi-organ failure (MOF) caused by one or more severe dysfunctions, such as: acute lung injury (ALI), adult respiratory distress syndrome (ARDS), hypermetabolism, hypotension, persistent inflammation, and / or sepsis. Renal function may also be impaired due to renal damage associated with the administration of nephrotoxic drugs as part of a surgery (such as angiography, diabetes, autoimmune diseases, and other dysfunctions and / or injuries associated with renal damage). To evaluate a patient's condition and long-term monitor the severity and / or progression of renal function, there is considerable interest in developing a simple, accurate, and continuous method for determining renal failure, preferably through a non-invasive procedure.

[0006] Serum creatinine concentration, an endogenous marker of renal function, is typically measured from a blood sample and used in combination with patient demographic factors such as weight, age to estimate glomerular filtration rate (GFR), which is a measure of renal function. However, due to many potential factors including: age, hydration status, renal perfusion, muscle mass, dietary intake, and many other anthropometric and clinical variables, creatinine-based assessment of renal function may be prone to inaccuracies. To compensate for these differences, a series of creatinine-based equations (recently extended to cystatin C) have been established that include factors such as sex and other relevant factors for estimating glomerular filtration rate (eGFR) based on serum creatinine measurement. However, these eGFR equations do not provide any way to compensate for most of the above sources of variability and thus have relatively poor precision. In addition, the eGFR method typically produces results that lag the true GFR by 72 hours.

[0007] Existing methods for measuring GFR have employed exogenous labeled compounds such as inulin, phthalein, 51 Cr-EDTA, Gd-DTPA, and 99m Tc-DTPA. Other endogenous markers such as labeled orthoiodohippurate or 99m Tc-MAG3 labeled with 123 I and 125 I have been used in other existing methods for assessing the tubular secretion process. However, the use of typical exogenous labeled compounds may be accompanied by various undesirable effects, including introducing radioactive material and / or ionizing radiation into the patient's body, and 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] In situations where the patient is specific and the environment may vary, the availability of a method for measuring renal excretion rate in real-time, accurately, and reproducibly using an exogenous marker would represent a significant improvement over any currently practiced method. In addition, a method that relies solely on the renal clearance of an exogenous chemical entity 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] According to one aspect of the present application, a sensor head is disclosed, comprising: a housing configured to be attached to a patient's body surface and including a bottom surface; at least two light sources configured to transmit light into a first region of the patient, the at least two light sources including a first blue LED light source and a second green LED light source, wherein the first blue LED light source transmits light at an excitation wavelength; and at least one light detector configured to detect light at an emission wavelength at a second region of the patient; wherein the housing encloses the at least two light sources and the at least one light detector.

[0010] According to another aspect of the present application, a sensor head is disclosed, comprising a housing configured to be attached to a patient's body surface, the housing including: (i) an inner housing having a sensor mount; (ii) at least two light sources configured to transmit light into a first region of the patient, wherein at least one light source transmits light at an excitation wavelength; and (iii) at least two light detectors configured to detect light at an emission wavelength at a second region of the patient, wherein the housing encloses the at least two light sources and the at least two light detectors, wherein the sensor mount includes: a first detection well having a first light detector mounted therein; a second detection well having a second light detector mounted therein; and a light source well having a first light source and a second light source mounted therein, and wherein the first detection well, the second detection well, and the light source well are configured to be optically isolated from each other to ensure that light from the light sources does not reach the light detectors without being coupled through the patient's body surface. Description of the Drawings

[0011] The patent or application document contains at least one drawing executed in color. Copies of this patent or patent application publication, including the color drawing(s), will be provided by the Office upon request and payment of the necessary fees.

[0012] The present disclosure will be better understood when consideration is given to its following detailed description, and features, aspects, and advantages will become apparent in addition to those set forth above. This detailed description refers to the following drawings, in which:

[0013] Figure 1 is a schematic diagram of a single-wavelength kidney monitoring device in one aspect;

[0014] Figure 2 is a schematic diagram of a dual-wavelength kidney monitoring system in one aspect;

[0015] Figure 3Is a graph summarizing the absorption, transmission, and emission spectra of various devices, materials, and compounds related to non-invasive monitoring of exogenous fluorescent agents defined within the optical wavelength range of approximately 430 nm to approximately 650 nm in vivo;

[0016] Figure 4 Is a graph summarizing the absorption spectra of oxyhemoglobin (HbO2) and deoxyhemoglobin (Hb) defined within the optical wavelength range of approximately 200 nm to approximately 650 nm;

[0017] Figure 5 Is a schematic diagram of the timing of an optical pulse cycle associated with data acquisition in a dual-wavelength renal monitoring system, where each optical pulse cycle sequentially includes optical pulses generated at an excitation wavelength and an emission wavelength;

[0018] Figure 6 Is a side view of a sensor head of a renal function monitoring system in one aspect;

[0019] Figure 7 Is Figure 6 The bottom view of the sensor head.

[0020] Figure 8 Is Figure 6 The internal top view of the sensor head, Figure 8 Showing the arrangement of various electrical components within the housing of the sensor head of a renal function monitoring system in one aspect;

[0021] Figure 9 Is Figure 8 An enlarged view of the internal view.

[0022] Figure 10 Is a schematic diagram of holes formed within the contact surface of a sensor head in a renal function monitoring system in one aspect;

[0023] Figure 11 Is a schematic diagram of light being synchronously detected by a photodetector of a sensor head in one aspect;

[0024] Figure 12 Is a schematic diagram of optical signal modulation and demodulation performed by a sensor head in one aspect;

[0025] Figure 13 Is a block diagram showing the subunits of a processing unit in one aspect;

[0026] Figure 14A Is a flowchart showing the steps of a global error mapping method for determining the parameters of a diffuse reflection correction equation in one aspect;

[0027] Figure 14B Is a flowchart showing the steps of a global error mapping method for determining the parameters of a diffuse reflection correction equation in a second aspect;

[0028] Figure 15A is a graph of representative intrinsic fluorescence measurements of the fluorescent agent (IF agent ) detected by a renal monitoring device obtained before and after injection of an exogenous fluorescent agent. The data subset selected for analysis to determine the correction factor is highlighted in orange.

[0029] Figure 15B is a graph of representative intrinsic fluorescence measurements of the fluorescent agent (IF agent ) detected by a renal monitoring device obtained before and after injection of an exogenous fluorescent agent. The data subset selected for analysis to determine the correction factor by fitting IF agent to the plasma derivative IF agent is highlighted in orange.

[0030] Figure 16 is a graph comparing the logarithmically transformed single-exponential curve fit of the corrected fluorescence signal measurement (log[Fit], black dashed line) and the corrected fluorescence signal measurement (IF Figure 15A and Figure 15B ) according to Figure 15A and Figure 15B on a portion of the selected analysis region from agent (red line).

[0031] Figure 17 is a map of a representative error surface summarizing the normalized root mean square error (RMSE, color scale), which is calculated for the difference between the linear fit of the logarithm of fluorescence over the range of the correction factors K ex and K em,filtered and the corrected fluorescence signal measurement, where the minimum RSME region is identified by the white arrow overlaying the map;

[0032] Figure 18 is a graph comparing the raw (F, blue line) and corrected (IF, red line) fluorescence signal measurements obtained before and after injection of an exogenous fluorescent agent.

[0033] Figure 19 is a flowchart summarizing the steps of a linear regression model method for determining the parameters of a diffuse reflectance correction equation in one aspect;

[0034] Figure 20 is a graph of the logarithmically transformed raw fluorescence signal (Log(Flr)), showing the regions of the data used as projections suitable for: a linear regression model for developing a data correction algorithm (orange line), the response variable for the linear regression model (black dashed line), and the region of data with significant variation for training the linear regression model (blue line);

[0035] Figure 21A is a graph of the raw fluorescence signal measurements obtained before and after injection of an exogenous fluorophore. The raw fluorescence signal measurements were obtained during exposure to various perturbations, represented as the colored regions, starting at approximately 13:50 hours. The various perturbations include changes in blood oxygen in the test subject, application and removal of pressure to the measurement area, administration of blood pressure medications to the test subject, cooling of the measurement area, and removal / replacement of the sensor head of the device;

[0036] Figure 21B is Figure 21A a graph of the corrected fluorescence signal measurements;

[0037] Figure 21C is Figure 21A a graph of the diffuse reflectance signal measurements taken simultaneously with the raw fluorescence signal measurements of Figure 21A These signals are used to correct the raw fluorescence signal measurements of Figure 21B to produce the corrected signal shown in

[0038] Figure 22A is a block diagram showing the multiple modules of a preprocessing sub-unit of one aspect;

[0039] Figure 22B is a block diagram showing the multiple modules of a preprocessing sub-unit of a second aspect;

[0040] Figure 23 is an isometric view of the sensor head of a second aspect renal function monitoring system;

[0041] Figure 24 is Figure 23 a bottom view of the sensor head of the renal function monitoring system shown in

[0042] Figure 25 is Figure 23 an isometric view of the sensor head of the renal function monitoring system shown in

[0043] Figure 26 is Figure 25 an exploded view of the inner housing of the sensor head shown in

[0044] This written description uses examples to disclose the invention, including the best mode, and also enables those skilled in the art to practice the invention, including making and using any device or system and performing any combined method. 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 materially 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, the preferred materials and methods are described below.

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

[0047] As used herein, a measurement value 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 set including demodulated in-phase, demodulated quadrature, and average measurement values from one acquisition channel.

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

[0050] As used herein, a measurement set refers to a set including one subset of measurements for each source LED.

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

[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 collected from a single channel of a telemetry ADC.

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

[0055] Figure 1FIG. 100 is a schematic diagram of system 100 provided as a non-limiting example, where a photodetector 110 configured to detect only those photons having an emission wavelength (λ em ) is used to detect fluorescence 102 having an emission wavelength (λ em ) from a region of interest of patient 104. Generally, an exogenous fluorophore 112 generates fluorescence 102 in response to an excitation event, which includes but is not limited to: illumination by light 106 at an excitation wavelength (λ ex ), the occurrence of an enzymatic reaction, a change in local electric potential, and any other known excitation event associated with the exogenous fluorophore. In one aspect, system 100 may include a light source 108 configured to deliver light 106 at an excitation wavelength (λ ex ) to patient 104. In this regard, fluorescence 102 is generated in response to illumination by light 106. Additionally, the excitation wavelength (λ ex ) of light 106 and the emission wavelength (λ em ) of fluorescence 102 are spectrally different (i.e., λ ex is completely different from λ em ), such that the photodetector 110 can be configured to selectively detect only fluorescence 102 by including any known light wavelength separation means, which includes but is not limited to an optical filter.

[0056] In some aspects, changes in fluorescence 102 can be monitored to obtain information about the physiological function or state of the patient. As a non-limiting example, the time-dependent decrease in fluorescence 102 measured after introducing the exogenous fluorophore 112 into the circulatory vessels of patient 104 can be analyzed to obtain information about the renal function of patient 104. In this non-limiting example, it can be assumed that the rate of decrease in fluorescence 102 is proportional to the rate at which the kidneys of patient 104 remove the exogenous fluorophore 112, thereby providing a measurement of renal function, including but not limited to: renal decay time constant (RDTC) and glomerular filtration rate (GFR).

[0057] Without being limited to any particular theory, the intensity of fluorescence 102 detected by the photodetector 110 can be affected by any one or more of a variety of factors, including but not limited to: the intensity or power of light 106 delivered to patient 104 at λ ex , the scattering and absorption of light 106 by the intervening tissue 114 of patient 104 between the light source 108 and the exogenous fluorophore 112, the concentration of the exogenous fluorophore 112 illuminated by light 106, and the scattering and absorption of fluorescence 102 by the intervening tissue 114 of patient 104 between the exogenous fluorophore 112 and the photodetector 110 at λ em .

[0058] Existing methods typically assume that the optical properties within the intervening tissue 114 remain substantially constant throughout the entire period during which the system 100 obtains measurements. As a result, existing methods typically obtain an initial measurement through the intervening tissue 114 of the patient 104 prior to the introduction of the exogenous fluorophore 112 and subtract these initial measurements to correct all subsequent data obtained after the introduction of the exogenous fluorophore 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 optical coupling efficiency of the light detector 110 to the patient 104; the concentration of chromophores such as hemoglobin due to changes in blood volume caused by vasodilation, vasoconstriction, or compression; changes in the optical properties of chromophores such as hemoglobin due to changes in the oxygenation state; and changes in the tissue structure such as changes associated with edema.

[0059] These dynamic changes in the optical properties of the intervening tissue 114 can introduce uncertainty into long-term measurements of fluorescence 102. As a non-limiting example, a change in the optical properties of the intervening tissue 114 can modulate the intensity or power of the light 106 that irradiates the exogenous fluorophore 112, resulting in modulation of the fluorescence 102 produced by the exogenous fluorophore 112, which can be misinterpreted as modulation of the concentration of the exogenous fluorophore 112. As another non-limiting example, a change in the optical properties of the intervening tissue 114 can modulate the intensity or power of the fluorescence 102 that reaches the light detector 110, which can also be misinterpreted as modulation of the concentration of the exogenous fluorophore 112. The potential modulation of changes in the optical properties of the intervening tissue 114 can introduce uncertainty into measurements of fluorescence 102, particularly those associated with long-term monitoring of fluorescence 102, as described above.

[0060] In various aspects, a method is provided for correcting in vivo real-time measurements of fluorescence from an exogenous fluorophore to eliminate the effects of changes in optical properties within patient tissue. An additional measurement of light passing through the patient tissue via an optical path that is separate from the optical path 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 an exogenous fluorophore within the patient. It has been found that including this additional measurement in the correction method significantly enhances the fidelity of the fluorescence measurement even in the presence of significant perturbations as described below.

[0061] A detailed description is provided below of an apparatus for monitoring fluorescence of an exogenous fluorophore in vivo and a method for correcting fluorescence measurements to eliminate the effects of diffuse reflection of light within patient tissue.

[0062] Although the devices and methods are described below in the context of a non-invasive optical renal function monitor, it should be understood that, with appropriate modifications, the calibration methods described herein can be applied to any compatible device configured to perform measurements by transmitting EM radiation from an external source and / or receiving EM radiation that has propagated through any scattering medium to an external detector via any scattering medium. 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 living or non-living material capable of propagating EM radiation at at least one EM frequency without limitation. 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 the scattering medium include: tissue of a living or dead organism, such as mammalian skin; gas, such as air with or without additional particles (such as dust, droplets, or solid particulate material); fluid, such as water with or without additional particles (such as bubbles or solid particulate material). Additionally, 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 system or gastrointestinal system.

[0063] System description

[0064] In various aspects, a method for calibrating fluorescence measurements to eliminate the effects of local skin property variations can be incorporated into any fluorescence monitoring system, including but not limited to, a system for optically monitoring renal function in vivo in real time by measuring the fluorescence changes of an exogenous fluorescent agent injected into a patient as the agent is eliminated from the patient's body. Figure 2 FIG. is a block diagram of a system 200 for optically monitoring the renal function of a patient 202 by fluorescence measurement of an exogenous fluorescent agent injected into the patient 202. The system 200 can include at least one sensor head 204 configured to transmit light at an excitation wavelength (λ ex ) into a first region 206 of the patient 202. The system 200 is also configured to detect light at an emission wavelength (λ em ) at a second region 208 of the patient 202, and to detect light at the excitation wavelength (λ ex ) and / or the emission wavelength (λ em ) at a third region 210 of the patient 202.

[0065] System 200 may also include a controller 212, an operation unit 214, and a display unit 216 that are operatively coupled to at least one sensor head 204. In various aspects, the controller 212 is configured to control the operation of at least one sensor head 204, as described in further detail below. The controller 212 is also configured to receive measurements of light from at least one sensor head 204. The controller 212 is further configured to correct the light measurements corresponding to fluorescence from an exogenous fluorophore according to at least one method, including but not limited to the method of using diffuse reflection of light to correct fluorescence measurements as disclosed. The controller 212 is also configured to convert the fluorescence measurements received from at least one sensor head 204 into summary parameters representative of the renal function of patient 202. In addition, the controller 212 is configured to receive at least one signal representative of user input from the operation unit 214 and generate one or more forms for display on the display unit 216, including but not limited to a graphical user interface (GUI).

[0066] A detailed description of the sensor head 204 and the controller 212 is provided below.

[0067] A. Sensor Head

[0068] In various aspects, the sensor head 204 includes at least one light source and at least one light detector in a housing. Figure 6 is a side view of a housing 600 for the sensor head 204 in one aspect, the housing 600 including an upper housing 602 and a lower housing 604 that are joined 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 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 may be fixed to the skin of patient 202. In various aspects, the adhesive material may 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 may be an optically transparent material. In another aspect, the adhesive material may be made of a non-fluorescent material to prevent the adhesive material from producing confounding fluorescence.

[0069] In various other aspects, the upper housing 602 may also include one or more openings 806 that are configured to provide access to the interior of a cable, including but not limited to a USB cable, and / or provide a window for a display (such as an indicator LED) generated by a circuit contained within the housing 600.

[0070] Figure 7 is Figure 8Bottom view of the shown housing 600. The contact surface 606 may include an aperture plate 702 that includes one or more apertures 704 configured to transmit light between the patient's skin and a light source and a light detector housed within the housing 600. In one aspect, the aperture plate 702 may 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 one or more of the apertures 704 may 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 may further 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.

[0071] Figure 8 is a schematic diagram showing the arrangement of the electrical components within the housing 600. Referring to Figure 8 , the upper housing 602 and the lower housing 604 may be fixed together with screws 802, and the screw holes and the interface between the two housing pieces may be filled with a waterproof filling material 804 including but not limited to a silicone material such as room temperature vulcanizing silicone (RTV) to prevent liquid from entering the interior of the housing 600.

[0072] In one aspect, the housing 600 may further include a cable opening 806 formed through the upper housing 602. The cable opening 806 may be configured to provide access to the interior of a cable including but not limited to a USB cable. In one aspect, the cable may be capable of powering a light source, a light detector, an indicator light, and associated electrical devices and circuits, as described below. In another aspect, the cable may also enable control signals to be transmitted into the housing to enable the operation of the electrical components within the housing 600, and the cable may also enable data signals to be transmitted that encode measurements obtained by one or more of the sensor devices included within the housing 600, the sensor devices including but not limited to: a first light detector 222, a second light detector 224, any additional light detectors such as a first monitor photodiode 904 and a second monitor diode 906, and any additional temperature sensors 228 (see Figure 9 ). In one aspect, the cable may 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 resin, and may also be sealed with a waterproof filling material including but not limited to RTV to prevent water intrusion.

[0073] In another aspect, the housing 600 may further include at least one display opening 808 formed through the upper housing 602. In one aspect, each display opening 808 may be configured to provide a window for a display (such as indicator LED 810) generated by circuitry contained within the housing 600. In one aspect, each indicator LED 810 may be positioned on a circuit board 812. In one aspect, a light pipe 814 may be epoxied into the display opening 808 within the upper housing 602 above each indicator LED 810. Each light pipe 814 may be filled with a waterproof filling material, such as RTV, for protection against liquid ingress. In various aspects, at least one indicator LED 810 may illuminate in a predetermined pattern so that a user of the system 200 can monitor the operating state of the sensor head 204.

[0074] Figure 9 is a close-up view of the internal optical region of the sensor head 204, showing in one aspect the arrangement of the light sources 218 / 220 and the 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 is separated from the second light detector 224 by a sensor mount 912 fixed 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 in the first detection well 908 and the second light detector 224 in the second detection well 910 ensures that the fluorescence signal generated by an exogenous fluorophore within the tissue of the patient 202 is distinguishable from the unfiltered excitation light introduced by the first light source 218.

[0075] Referring again to Figure 9 , the sensor mount 912 may 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 may be held in place using screws 916. In one aspect, the sensor mount 912 may be fixed 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 apertures 704 configured to transmit light to and from the skin beneath the contact surface 606 of the sensor head 204 are formed through a structurally separate aperture plate 702 (see Figure 7) is formed to provide precise alignment of the aperture 704 with the corresponding light sources 218 / 220 and the light detectors 222 / 224, which will be described in more detail below.

[0076] 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 ground of the circuit board using conductive screws 916. Additionally, any glass windows (including but not limited to the optical filter 244 and the transparent glass 246 as described below (see Figure 2 )) located within the source well 902 and / or the detector wells 908 / 910 adjacent to the aperture plate 702 can further include a conductive coating. Non-limiting examples of suitable conductive coatings for the glass windows of the sensor mount include conductive indium tin oxide (ITO) coatings and any other suitable transparent and conductive coatings.

[0077] Without being limited to any particular theory, the conductive material of the sensor mount 912 provides a partial Faraday cage to shield the electrically sensitive detectors 222 / 224 from electrical noise generated or conducted by the patient's body. The partial Faraday cage provided by the sensor mount 912 can complete the conductive ITO coating on the glass windows within the source well 902 and / or the detector wells 908 / 910. In one aspect, the conductive coating on the glass window, such as the ITO coating, is sufficiently conductive to provide electrical shielding while maintaining sufficient transparency to transmit light to and from the skin surface of the patient 202. In another aspect, the ITO coating of each glass window can be grounded to the conductive sensor mount 912 using any known electrical grounding method, including but not limited to: a wire connecting the glass coating to the sensor mount 912, which is attached at both ends of the wire 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.

[0078] 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 transparent double-sided medical-grade adhesive, as Figure 6 and Figure 7As shown. As described herein, any adhesive material is selected to be optically transmissive at the excitation and emission wavelengths used by system 100. The adhesive material 610 may be positioned on the contact surface 606 such that the adhesive material covers the holes 704, but exposes the temperature sensor openings 706 to ensure sufficient thermal contact with the skin of patient 202. In one aspect, the sensor head 204 may be further secured to the patient 202 using one or more additional biocompatible medical fastener devices as needed, including but not limited to: Tegaderm bandages, medical tapes, or any other suitable biocompatible medical fastener device).

[0079] In one aspect, the contact surface 606 may be located near the leading edge of the sensor head 204 to provide accurate positioning of the contact surface 606 over a selected area of the patient's skin. In another aspect, the holes 704 may be positioned toward the center of the contact surface 606 to reduce ambient light entry. Without being limited to any particular theory, ambient light may enter one or more of the holes 704 due to incomplete adhesion of the contact surface 606 to the patient's skin and / or due to ambient light passing through the patient's exposed skin just outside the covered area of the contact surface 606 and into the holes 704.

[0080] Referring again to 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 attach 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 may be filled with a biocompatible foam to ensure consistent contact with the patient 202.

[0081] i) Light source

[0082] 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 the 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 may be selected to fall within a spectral range where the exogenous fluorophore exhibits relatively high absorption. In another aspect, the emission wavelength may be selected to fall within a spectral range where the exogenous fluorophore exhibits relatively high emission. The exogenous fluorophore may 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 fluorophore may be selected to be measured within a spectral range where absorption changes of other chromophores such as hemoglobin within the tissue of the patient 202 are relatively low during use.

[0083] Without being limited to any particular theory, hemoglobin (Hb) is an absorber of visible light in the tissue of patient 202, and if the Hb absorber changes during the measurement period of system 200, it is possible to interfere with the fluorescence measurement of the exogenous fluorophore. Since hemoglobin (Hb) can carry out gas exchange in almost all tissues containing circulating blood vessels, almost all tissues are vulnerable to interference with the fluorescence measurement of system 200 due to fluctuations in hemoglobin concentration. In most tissues, externally applied pressure may cause blood pooling, which may be manifested as an obvious attenuation of the fluorescence measured on the skin surface. The periodic opening and closing of blood vessels near the skin surface ("vasomotion") may also cause fluctuations in hemoglobin concentration, which may introduce additional noise to the fluorescence measurement of the exogenous fluorophore by system 200. In addition, in some patients 202, such as those suffering from lung diseases, changes in the Hb oxygenation state can also be observed, resulting in additional potential changes in background skin absorption due to the difference in the absorption spectra of deoxyhemoglobin (Hb) and oxyhemoglobin (HbO2), as Figure 3 shown.

[0084] In one aspect, the excitation wavelength and emission wavelength for the exogenous fluorophore can be selected to coincide with a pair of HbO2 / Hb isosbestic points, and each isosbestic point defined herein is a wavelength characterized by approximately equal light absorbance of HbO2 and Hb. Without being limited to any particular theory, as long as the combined concentration of HbO2 and Hb remains relatively stable during the fluorescence measurement by system 200, the fluorescence measurement performed at each isosbestic wavelength is less sensitive to changes caused by changes in the oxygenation of hemoglobin. 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.

[0085] In various aspects, the excitation wavelength and emission wavelength can be selected based on the absorption wavelength and emission wavelength of the selected exogenous fluorophore of system 200. In one aspect, the excitation wavelength can be an HbO2 / Hb isosbestic wavelength and can simultaneously be a wavelength within the high absorbance spectral range of the exogenous fluorophore. In another aspect, the emission wavelength can be an HbO2 / Hb isosbestic wavelength and can simultaneously be a wavelength within the spectral range emitted by the exogenous fluorophore. Table 3 provides a summary of HbO2 / Hb isosbestic wavelengths in the spectral range from 200 nm to approximately 1000 nm. Figure 4 is a graph of the absorption spectrum for identifying the HbO2 / Hb isosbestic wavelengths in Table 1.

[0086] Table 1. Absorbance Wavelengths of HbO2 / Hb etc., λ = 200 to 1000 nm

[0087]

[0088]

[0089] As an illustrative example, Figure 3 is a graph that summarizes, in one aspect, the absorption spectra of HbO2 and Hb, as well as the absorption and emission spectra of the frequency spectrum of the exogenous fluorophore MB-102. The emission spectra of the blue LED light source and the green LED light source are also shown superimposed on Figure 3 the other spectra. In this aspect, system 200 can include a blue LED as the first light source 218, and the excitation wavelength of system 200 can be an absorbance wavelength of approximately 450 nm. As listed in Table 1 and Figure 3 shown, the Hb absorbance spectrum slopes significantly at absorbance wavelengths from 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 an absorbance 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 slopes less steeply, and a broader bandwidth light source (including but not limited to an LED with a bandpass filter) may be sufficient to be used as the first light source 218.

[0090] In another aspect, the excitation wavelength can be selected to enhance the light absorption contrast between the exogenous fluorophore and the chromophore within the tissue of patient 202. As a non-limiting example, as Figure 3 shown, at an absorbance wavelength of 452 nm, the light absorption of MB-102 is more than three times higher than the light absorption of HbO2 and Hb. Without being limited to any particular theory, relative to HbO2 and Hb, MB-102 will absorb a higher proportion of the light that irradiates the tissue of patient 202 at a wavelength of approximately 450 nm, thereby enhancing the absorption efficiency of MB-102 and reducing the intensity of the light at the excitation wavelength, which is required to initiate a detectable fluorescence signal.

[0091] In various aspects, a second absorbance wavelength can also be selected as the emission wavelength of 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 chromophores at the selected emission wavelength can reduce the loss of light emitted by the exogenous fluorophore and improve the efficiency of fluorescence detection.

[0092] In various aspects, the first light source 218 and the second light source 220 can be any light sources configured to transmit light at an excitation wavelength and an emission wavelength. Generally, the first light source 218 transmits light to the exogenous fluorophore at an intensity sufficient to penetrate the tissue of the patient 202, with sufficient remaining intensity to induce the exogenous fluorophore to emit light at the emission wavelength. Generally, the first light source 218 transmits light to the exogenous fluorophore at an intensity sufficient to penetrate the tissue of the patient 202, where the intensity remaining after scattering and / or absorption is sufficient to induce fluorescence at the emission wavelength by the exogenous fluorophore. However, the intensity of the light transmitted by the first light source 218 is limited to an upper limit value to prevent adverse effects on the exogenous fluorophore and / or endogenous chromophores ("autofluorescence") in the skin, such as tissue burning, cell damage, and / or photobleaching.

[0093] Similarly, the second light source 220 transmits light at the emission wavelength of the exogenous fluorophore at an intensity configured to provide sufficient energy to propagate through the first region 206 of the patient by scattering and absorption and to propagate out of the second region 208 and the third region 210 with sufficient remaining intensity for detection by the first light detector 222 and the second light detector 224, respectively. Like the first light source 218, the intensity of the light generated by the second light source 220 is limited to an upper limit value to prevent adverse effects such as the tissue damage or photobleaching described above.

[0094] In various aspects, the first light source 218 and the second light source 220 can be any light sources suitable for use with fluorescence medical imaging systems and devices. Non-limiting examples of suitable light sources include: LEDs, diode lasers, pulsed lasers, continuous wave 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 that is suitable for monitoring the concentration of the exogenous fluorophore 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.

[0095] In one aspect, the selection of the intensity of the 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 allowable exposure (MPE) for skin exposure to a laser beam according to applicable regulatory standards such as ANSI standard Z136.1. In another aspect, the light intensity for the system 200 can be selected to reduce the likelihood of photo-bleaching of exogenous fluorescent sources and / or other chromophores within the tissue of the patient 202, which includes but is not limited to: collagen, keratin, elastin, hemoglobin within red blood cells, and / or melanin within melanocytes. In yet another aspect, the light intensity for the system 200 can be selected to elicit a detectable fluorescence signal from exogenous fluorescent sources within the tissue of the patient 202 and / or within the first light detector 222 and / or the second light detector. In yet another aspect, the light intensity for the system 200 can be selected to provide an appropriately high light energy while reducing power consumption, suppressing heating / overheating of the first light source 218 and the second light source 220, and / or reducing the time that the patient's skin is exposed to light from the first light detector 222 and / or the second light detector.

[0096] In various aspects, the intensity of the first light source 218 and the second light source 220 can be modulated to compensate for any one or more of at least several factors, which at least several factors include but are not limited to: individual differences in chromophore concentration within the 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 characteristics. In one aspect, variations in skin pigmentation can be between two different individual patients 202, or between two different locations on the same patient 202. In one aspect, the light modulation can compensate for variations in the optical path taken by the light passing through the tissue of the patient 202. The optical path can vary due to any one or more of at least several factors, which at least several factors include but are not limited to: variations in the separation distance between the light source and the light detector of the system 200; variations in the secure attachment of the sensor head 204 to the skin of the patient 202; variations in the light output of the light source due to the light source being exposed to environmental factors such as heat and humidity; variations in the sensitivity of the light detector due to the light detector being exposed to environmental factors such as heat and humidity; modulation of the duration of illumination by the light source, and any other relevant operating parameters.

[0097] In various aspects, the first light source 218 and the second light source 220 can be configured to modulate the intensity of the light generated as needed according to 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 vary the output flux as needed, such as LED light sources, methods including but not limited to modulating the electrical potential, current, and / or power supplied to the first light source 218 and / or the second light source 220 can be used to electronically modulate the intensity of the light. In another aspect, optical methods can be used to modulate the intensity of the light, and the optical methods include but are 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 the light leaving the first light source 218 and the second light source 220; using optical devices including but not limited to lenses, mirrors, and / or prisms to deflect the path of the light leaving the first light source 218 and the second light source 220 away from the first region 206 of the patient.

[0098] 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 rate of energy within the generated light beam. In one aspect, the laser flux can be limited to a range defined by safety standards, which include but are 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, which include but are not limited to the wavelength of the light being 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 in the wavelength range of about 1500 nm to about 1800 nm, with a duration up to about 10 seconds. For light delivered in the wavelength range of about 400 nm to about 1400 nm (visible light / NIR light), the maximum flux can be about 0.6 J / cm2, with a duration up to about 10 seconds, and about 0.2 J / cm2, with a duration in the range of about 10 seconds to about 30000 seconds. For extended exposure, according to the ANSI standard, the light delivered is limited to a maximum power density (W / cm2): visible light / 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, according to the ANSI standard, extended exposure to light delivered at UV wavelengths is generally not recommended.

[0099] 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 fluorophore without photobleaching, and to irradiate the exogenous fluorophore 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, and 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 fluxes generated by the light source at 450 nm or 500 nm can be limited to 1.5 mW / cm 2 and 5 mW / cm 2 , respectively, to prevent photobleaching.

[0100] In various aspects, in the absence of the limitations described above herein, the light fluxes generated by the first light source 218 and the second light source 220 can be modulated by any suitable system and / or device. This modulation can be enabled once during the operation of the system 200, and thus, the light fluxes generated by each of the first light source 218 and the second light source 220 can be relatively constant throughout the operation of the system 200. In another aspect, this light modulation can be enabled at discrete times during the operation of the system 200, or this light modulation can be continuously enabled during the operation of the system 200.

[0101] In one aspect, when the system 200 is configured in engineering mode, the light flux can be modulated via manual adjustment of any of the power settings and / or optical device settings 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, the sensors including but not limited to additional light detectors 226 and temperature sensors 228, as described in more detail below.

[0102] 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, which is defined herein as the duration of the generated light. While pulse width is typically used to characterize the performance of light sources that generate light in discrete pulses (such as pulsed lasers), 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, such that a single fluorescence measurement can be obtained through 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: exogenous fluorophores or other chromophores within the tissue of the patient 202 without photobleaching, delivering sufficient light energy to elicit detectable fluorescence from the exogenous fluorophore; complying with safety standards such as ANSI standards for the light delivered to the patient; delivering light at a high enough rate to be able to acquire data at a rate compatible with real-time monitoring of renal function; the performance capabilities of the selected light sources, light detectors, and other devices of the system 200; preservation of the service life of the light sources, light detectors, and other devices related to generating and detecting light energy; and other relevant factors.

[0103] In various aspects, the pulse width of the light generated by the first light source 218 and the second light source 220 can be independently selected to be a duration in the range from about 0.0001 seconds to about 0.5 seconds. In various other aspects, the pulse width of the light generated by the first light source 218 and the second light source 220 can be independently selected to be a duration in the range 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 width of the light generated by the first light source 218 and the second light source 220 is both about 0.1 seconds, as Figure 5 schematically shown in.

[0104] 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 the pulse rate is typically used to characterize the performance of light sources that produce discrete pulsed light (such as pulsed lasers), it should be understood that the term "pulse rate" as used herein refers to the rate at which a single light source produces discrete light pulses at a single wavelength, which rate 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 a patient; the performance capabilities of the selected light sources, light detectors, and other devices of the system 200; the rate of delivering light being compatible with a data acquisition rate that is fast enough for real-time monitoring of renal function; preservation of the service life of the light sources, light detectors, and other devices related to the generation and detection of light energy; and any other relevant factors.

[0105] In various aspects, the light source is configured to deliver light to the tissue of the patient 202 at a single location (such as the first region 206), as Figure 2 schematically illustrated. In one aspect, delivering light of both the excitation wavelength and the emission wavelength to the same first region 206 enables two light pulses to share at least a portion of the optical path through the tissue of the patient 202 between the entry point at the first region 206 and the detection points at the second region 208 and the third region 210. As discussed in detail below, this arrangement of the optical path enhances the quality of the data generated by the system 200.

[0106] In one aspect, the first light source 218 and the second light source 220 are operatively coupled to a common device for light delivery. In one aspect (not shown), each of the first light source 218 and the second light source 220 can be operatively 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, which 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 are operatively coupled to a common optical fiber or other optical component, which 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.

[0107] In one aspect, system 200 may include a sensor head 204 that is provided with a sensor mount 912 configured with one or more wells in which a light source 218 / 220 and a light detector 222 / 224 may be attached in a predetermined arrangement. In one aspect, as Figure 9 and Figure 10 shown, a first light source 218 and a second light source 220 may be located within a source well 902 of the sensor mount 912 positioned within the sensor head 204 (see Figure 9 ). In one aspect, the source well 902 may contain a first LED light source 218 that produces light at an excitation wavelength and a second LED light source 220 that produces light at an emission wavelength, and the first LED light source 218 and the second LED light source 220 may be operatively coupled to a single light transmission aperture 1002 formed through a perforated plate 702 (see Figure 10 ), and the single light transmission aperture ensures that light of two wavelengths (i.e., excitation and emission) enters the skin of the patient 202 at substantially the same location (including but not limited to a first region 206), as Figure 2 schematically shown. In one aspect, the source well 902 also contains a first monitor photodiode 904 and a second monitor photodiode 906 that are used to correct for variations in the output power from the LED light sources, as described in further detail below.

[0108] In one aspect, only a portion of the light energy produced by the LED light sources is transmitted via the single light transmission aperture 1002 to the skin of the patient 202. In one aspect, the skin of the patient 202 receives approximately 1% of the light energy produced 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 produced by the LED light sources. Without being limited to any particular theory, a portion of the light produced by the LED light sources transmitted to the skin of the patient 202 may be increased by incorporating additional optical elements configured to focus and / or direct the light from each LED light source to the light transmission aperture 1002. In another aspect, a diffuser may be used to mix the output of the light sources such that the light energy is presented uniformly at the surface of the patient's skin.

[0109] ii) Light detector

[0110] Referring again to Figure 2, in various aspects, the system 200 further includes 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 includes an optical filter 244 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. With the tissue of the patient 202 irradiated with light at only the excitation wavelength and only 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 fluorophore and correct the fluorescence measurement by removing the effect of diffuse reflection of light according to the correction method described below.

[0111] In various aspects, light from the second region 208 and the third region 210 within the tissue of the patient 202 is detected by the first light detector 222 and the second light detector 224, respectively, and the second region 208 and the third region 210 are each nominally spaced apart from the first region 206. The light generated by the first light source 218 and the second light source 220 is transmitted to the first region 206. The nominal spacing distance can be selected to balance two or more effects that may affect the quality of the data detected by the light detector. Without being limited to any particular theory, as the nominal spacing distance increases, the total detection signal from the light detector 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 the emission wavelength, which can result in a less significant decrease in the detected fluorescence signal (i.e., the light at the emission wavelength) relative to the signal associated with the light detected at the excitation wavelength as the nominal spacing distance increases. The longer the nominal spacing distance, the higher the sensitivity to signal variations due to changes in the optical properties of the tissue.

[0112] In one aspect, the nominal spacing distance can range from 0 mm (i.e., co - location of the light source and the light detector) to about 10 mm. In various other aspects, the nominal spacing distance can range 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 spacing 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 spacing distance can be about 4 mm to balance these competing effects of the logarithmic decrease in the signal and the reduced size of the background signal relative to the signal from the exogenous fluorophore.

[0113] Referring again to Figure 9 , the first light detector 222 may be positioned within the first detection well 908 of the sensor mount 912, and the second light detector 224 may 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 may 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 delivery aperture 1002 are separated from each other by a nominal spacing distance disclosed hereinabove, including but not limited to a nominal spacing 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 may be optically isolated from each other to ensure 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. As described in detail below, the separation between the two detection wells 908 / 910 ensures that the detected fluorescence signal from the exogenous fluorophore can be distinguished from the unfiltered excitation light.

[0114] In one aspect, the three apertures 704 of the aperture plate 702 (see Figure 7 ) are circular and have a diameter range from about 0.5 mm to about 5 mm. In various other aspects, the diameter of the aperture may be in the range of 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.

[0115] In one aspect, the three apertures 704 of the aperture plate 702 are circular apertures having a diameter of about 1 mm. Due to the logarithmic decrease in the signal as the spacing distance from the light source at the skin interface of the sensor head 204 increases, this limited width of the aperture may result in an effective source-detector separation that is less than the nominal spacing distance.

[0116] In various aspects, the photodetectors 222 / 224 of system 200 can be any suitable light detection device, but are not limited thereto. Non-limiting examples of suitable light detection devices include: photoemissive 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 detection device. In one aspect, the photodetectors 222 / 224 are sensitive enough to detect fluorescence emitted by an exogenous fluorophore within the tissue of patient 202, the tissue of patient 202 including melanin in the range of about 1% to about 40% in the epidermis and a blood volume of about 0.5% to about 2% of the skin volume. In one aspect, the photodetectors 222 / 224 can be silicon photomultiplier (SPM) devices.

[0117] In one aspect, the first photodetector 222 can be configured to detect light at both the excitation frequency and the emission frequency, and the second photodetector 224 can be configured to detect only light at the emission frequency. In one aspect, due to the design and material of the sensor element of the second photodetector 224, the second photodetector 224 can respond only to light of the emission wavelength. In another aspect, the second photodetector 224 can respond to a wider range of light wavelengths, but can be located downstream of an optical filter configured to pass only a portion of the incident light having the emission wavelength and further configured to block the passage of light of wavelengths outside the emission wavelength.

[0118] Any suitable optical filter can be selected for use with the second photodetector 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 an absorption filter does not vary significantly with the angle of the incident light, whereas the performance of an interference / dichroic filter 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 patient 202.

[0119] In one aspect, the second photodetector 224 can be located downstream of an absorptive long-pass filter configured to pass light having a wavelength greater than a predetermined wavelength to the second photodetector 224. As a non-limiting example, the second photodetector 224 can be located downstream of a long-pass OG 530 filter configured to pass light having a wavelength greater than about 530 nm. Other non-limiting examples of suitable filters include the Hoya O54 filter and the Hoya CM500 filter.

[0120] In various aspects, an absorption filter 244 configured to absorb light at an excitation wavelength can be located within a second detection well 910 between a second photodetector 224 and a second detector aperture 1006. In one aspect, the absorption filter 244 can be constructed of OG 530 Schott glass. The thickness of the absorption filter 244 can be selected such that the optical density is sufficient to filter the excitation light by an amplitude of approximately three orders of magnitude. In one aspect, the thickness of the absorption filter 244 can range from about 1 mm to about 10 mm. In various other aspects, the thickness of the absorption filter 244 can range 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 thickness of the absorption 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 absorption filter 244 is a 3-mm thick filter constructed of OG 530 Schott glass.

[0121] In another aspect, an optical diffuser can be disposed within the light source well 902. In this aspect, the optical diffuser is capable of mixing the light entering the light source well 902 from the first light source 218 / 220 and the second light source 218 / 220. 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 irradiating 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 the non-mixed light, thereby reducing potential sources of variation.

[0122] In one aspect, a transparent material configured to pass light at both the excitation wavelength and the emission wavelength can be located within a first detection well 908 between a first photodetector 222 and a first detector aperture 1004. In this regard, the transparent material can be any material having optical properties similar to those of the material of the absorption filter 244, the optical properties 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 having the same thickness as the absorption filter 244.

[0123] As a non-limiting example, in Figure 3 the transmission spectrum of an OG 530 filter is provided. As Figure 3 shown, the transmission spectrum of the OG 530 filter overlaps with the emission spectrum of the MB-102 exogenous fluorophore and the emission spectrum (emission wavelength) of the green LED used as the second light source 220. Additionally, 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 fluorophore.

[0124] In one aspect, transparent materials such as glass 246 and absorption filter 244 can be respectively fixed to the flanges formed within the first detection well 908 and the second detection well 910. Transparent materials such as glass 246 and optical filter 244 can be fixed in place using an opaque and / or light-absorbing adhesive, which includes but is not limited to black epoxy resin, to ensure that all light received through the first detector aperture 1004 and the second detector aperture 1006 passes through the optical filter 244 or glass 246 before being detected by the first and second light detectors 222 / 224. In another aspect, the sides of the filter 244 or glass 246 can be painted black with a light-absorbing coating, which includes but is not limited to India ink, to ensure that light does not reach the first and second light detectors 222 / 224 without passing through the optical filter 244 or glass 246.

[0125] In one aspect, due to the Lambertian distribution of the angle of the light leaving the patient's skin, the height of the detection wells 908 / 910, in combination with the diameter of the detector apertures 1004 / 1006, can limit the proportion of light emitted from the second region 208 and the third region 210 of the patient's skin that reaches the effective area of the light detectors 222 / 224. In one aspect, the proportion of light emitted from the second region 208 and the third region 210 of the patient's skin received by the light detectors 222 / 224 can be in the range of approximately 5% to approximately 90%. In various other aspects, the proportion of light can be in the ranges of approximately 5% to approximately 15%, approximately 10% to approximately 20%, approximately 15% to approximately 25%, approximately 20% to approximately 30%, approximately 25% to approximately 35%, approximately 30% to approximately 40%, approximately 35% to approximately 45%, approximately 40% to approximately 60%, approximately 50% to approximately 70%, and approximately 60% to approximately 90%.

[0126] In one aspect, for Figure 6 and Figure 7 the sensor head 204 shown with holes 1002 / 1004 / 1006 having a diameter of 1 mm, approximately 10% of the light emitted from the patient's skin surface can reach the effective area of the light detectors 222 / 224 to be detected. In various aspects, the sensor head 204 can further include additional optical elements, which include but are not limited to lenses and / or prisms, configured to compensate for the Lambertian distribution of the light angle so as to enhance the proportion of light emitted from the patient's skin that is directed to the effective area of the light detectors 222 / 224.

[0127] iii) Temperature sensor

[0128] Referring to Figure 2, the 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 area. 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 the control methods for one or more temperature-sensitive devices of the system 200 described below.

[0129] As a non-limiting example, temperature measurement can 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. 11A ) can be used in the control scheme. Figure 11 ) to modulate the amount of power provided to the LED light source to compensate for the effect of the LED temperature on the LED light output. In another aspect, the additional temperature sensor 228 can monitor the temperature of the LED light source 218 / 220 to monitor and / or compensate for temperature changes of the LED, and monitor and / or compensate for the temperature-dependent transmission of the optical filter to maintain a relatively constant output wavelength.

[0130] As another non-limiting example, the additional temperature sensor 228 may be a thermistor 816 (see Figure 8 ) is included in the sensor head 204, and the thermistor 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, thermistor 816 can be epoxied into temperature sensor opening 706 in aperture plate 702. In this aspect, space 918 between circuit board (not shown) and lower housing 604 can be filled with thermally conductive putty to ensure good thermal conduction and heat dissipation.

[0131] 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, such that the LED light sources 218 / 220 can maintain a relatively constant output wavelength.

[0132] In another aspect, if an over-temperature condition is detected, the temperature measured by 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 photodetectors 222 / 224. In one aspect, if the housing temperature detected by the thermistor 816 is greater than about 40 °C, an over-temperature condition can be indicated. In various other aspects, an over-temperature condition where the housing temperature is greater than about 40.5 °C or greater than about 41.0 °C can be detected.

[0133] B. Controller

[0134] Referring again to Figure 2 , in various aspects, the system 200 can include a controller 212 that is configured to operate the light sources 218 / 200 and the photodetectors 222 / 224 in a coordinated manner to obtain a plurality of measurements for acquiring fluorescence of an exogenous fluorophore within the tissue of the patient 202, to correct the fluorescence data to remove the effects of diffuse reflection of light as described below, and to convert the fluorescence measurements into a parameter representative of the renal function of the patient 202. Figure 11 is a schematic diagram of the electronic circuit 1100 that, in one aspect, shows the arrangement of the various electrical components capable of operating the system 200. In one aspect, the controller 212 can be a computing device that further includes an operating unit 214 and a display unit 216.

[0135] i) Light source control unit

[0136] Referring again to Figure 2 , the controller 212 can include a light source control unit 230 that is configured to operate the first light source 218 and the second light source 220 to generate light at an excitation wavelength and an emission wavelength, respectively, in a cooperative manner to produce a repetitive pulse sequence, as Figure 5Schematically shown. 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, which include but are not limited to: activation or deactivation of each light source; relative timing of activation and deactivation of each light source to enable 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 specific light source 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 changes 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: the light output of the light sources 218 / 220 measured by the first monitor photodiode 904 and the second monitor photodiode 906 respectively within the source well 902, the temperature of the light sources 218 / 220, and any other feedback measurements related to monitoring the performance of the light sources 218 / 220.

[0137] 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 supplied 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, and the 16-bit DAC 1124 is operatively coupled to the LED current source 1126, as Figure 11 shown. In one aspect, the 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 supplied to the LED light sources 218 / 220 can be adjusted based on the waveform signal provided by the waveform generator / FPGA 1122.

[0138] Referring to Figure 5 , in one aspect, each light pulse sequence 500 includes emission wavelength light pulses 502 and excitation wavelength light pulses 504, both of which are composed of a plurality of square waves 506 generated by the first and second LED light sources 218 / 220. Referring to Figure 11 , the square waves generated by the waveform generator 1122 are received by the LED current source 1126. The current generated by the LED current source includes square waves similar to the waveform generated by the waveform generator 1122. Without being limited to any particular theory, since the intensity of the light generated by the LED light sources 218 / 220 is proportional to the amplitude of the received current, the light generated by the LED light sources 218 / 220 also includes as Figure 5The square wave shown. In another aspect, in the detailed discussion below, the square wave generated by the waveform generator 1122 can also be used by the acquisition unit 234 in a synchronous detection method to reduce the influence of various confounding factors, which include but are not limited to the detection of ambient light by the detector signals generated by the photodetectors 222 / 224 during the irradiation of the patient tissue by the first light source 218 / 220 and the second light source 218 / 220 at the emission wavelength and the excitation wavelength, respectively.

[0139] In various other aspects, various alternative LED pulse modulation schemes can be equivalently adopted without limitation. In one aspect, the excitation pulses and the emission pulses are transmitted in alternating series, with dark periods interspersed after each pulse. In another aspect, the first and second LED light sources 218 / 220 are each modulated with a 50% duty cycle but at different modulation frequencies, thereby allowing the signals associated with the excitation pulses and the emission pulses to be separated by frequency filtering.

[0140] Without being limited to any particular theory, the total optical power transmitted to the patient's skin can be limited by at least two factors: the photobleaching of exogenous fluorophores and / or endogenous chromophores, and the overheating of the patient tissue irradiated by the system 200. In one aspect, based on safety standards including but not limited to ANSI / IESNA RP-27.1-05, tissue heating can impose an absolute limit of approximately 9 mW on the optical power that can be transmitted to the skin. In another aspect, the photobleaching of the skin autofluorescence associated with endogenous chromophores (including but not limited to collagen, hemoglobin, and melanin) can provide a background signal to the measured fluorescence, which remains relatively constant as long as the chromophores are not self-bleached. This constant autofluorescence background can be subtracted from the original fluorescence signal, but if the autofluorescence changes 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 the power threshold associated with chromophore photobleaching.

[0141] Referring again to Figure 9 , in various aspects, the optical output of the light sources 218 / 220 can be measured using the monitor photodiodes 904 / 906. Since the light intensity reaching these monitor photodiodes 904 / 906 is generally much stronger than the light intensity reaching the photodetectors 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.

[0142] In various aspects, system 200 can be configured to operate within the 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 in amplitude. Additionally, due to the elimination of the reagent in the kidneys over time, the concentration of the exogenous fluorophore within each patient 202 can vary over a range of approximately two orders of magnitude in amplitude. In various aspects, system 200 can be configured to detect fluorescence from endogenous fluorophores over an intensity range of more than five orders of magnitude. In these various aspects, system 200 can be configured by modulating at least one operating parameter, which includes but is not limited to: the amplitude of the light output of light sources 218 / 220 and the sensitivity of photodetectors 222 / 224 corresponding to the detector gain.

[0143] In one aspect, the intensity of the light output by light sources 218 / 220 can be manually set by the 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 light sources 218 / 220. In one aspect, the light source control unit 230 can be configured to control the light intensity generated by LED light sources 218 / 220 within a range of normalized output intensities from 0 (off) to 1 (maximum power). In one aspect, the intensity of light sources 218 / 220 can be set by the light source control unit 230 in coordination with the detector gain of photodetectors 222 / 224 set by the photodetector control unit 232, as described below.

[0144] In one aspect, the light source control unit 230 can use the signals obtained during the first 10 detection cycles after data acquisition initialization but before injection of the exogenous fluorescent agent by the system 200 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 cycle 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 the light received by the light detectors 222 / 224 at the excitation and emission wavelengths during a detection cycle, the corresponding LED intensity can be modulated such that the analog signal generated by the light detectors 222 / 224 corresponds to approximately 1 / 4 of the full range of each detector 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 the light generated by the second LED light source 220 at the emission wavelength, the larger signal can be used to modulate the power setting of the second LED light source 220. If the above method results in an LED intensity setting modulated to 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 detectors 222 / 224 (i.e., 1 / 4 of the ADC range) is selected to retain additional light detection capabilities 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 plurality of factors, including but not limited to the introduction of the exogenous fluorescent agent into the patient 202.

[0145] In the above aspect, once the light source control unit 230 coordinates the setting of the LED intensity with the detector gain of the light detectors 222 / 224 set by the light detector control unit 232 within the first 10 detection cycles, an additional 10 detection cycles are obtained to confirm the applicability of these settings for the operation of the system 200 given the tissue properties of a particular patient 202, and then the LED intensity settings and detector gains are recalculated as described herein. If the newly calculated LED intensity is within twice the previously determined setting and the detector gain does not change, the previously determined setting is maintained for subsequent data acquisition cycles for determining renal function. Otherwise, the settings are updated using the same method described herein, and an additional 10 data acquisition cycles are performed to confirm the stability of the settings. 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.

[0146] ii) Light detector control unit

[0147] Referring again to Figure 2 , the controller 212 may include a photodetector control unit 232 configured to operate a first photodetector 222 and a second photodetector 224 to be able to detect light at the emission wavelength and unfiltered light at all wavelengths, respectively. In various aspects, the photodetector 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 photodetector control unit 232 may generate a plurality of optical measurement signals encoding the intensity of the light detected by the photodetectors 222 / 224, the optical measurement signals including, but not limited to, raw detector signals that may be received by an analog-to-digital converter (ADC) 1102 (see Figure 11 ). In another aspect, when the system 200 is configured in engineering mode, the detector gain and / or other detector control signals may be manually set by the user detector gain.

[0148] In various other aspects, the amount of light received by the photodetectors 222 / 224 may 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 fluorophore concentration within each patient 202, and any other relevant parameters. In one aspect, the gain of the first photodetector 222 and the second photodetector 224 may be set by the user via the operation unit 214. In another aspect, the photodetector control unit 232 may be configured to automatically modulate the gain of the photodetectors 222 / 224 via the bias voltage gain of a bias voltage generator 1112 (see Figure 11 ).

[0149] In one aspect, the signals obtained during the first 10 detection cycles obtained by the system 200 after data acquisition initialization but before injection of the exogenous fluorophore are used by the photodetector control unit 232 to automatically adjust the gain of the photodetectors 222 / 224 and the output intensity of the light sources 218 / 220. As previously described herein, an initial detection cycle may 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 may be modulated such that the analog signal generated by the photodetectors 222 / 224 corresponds to approximately 1 / 4 of the entire range of each detector analog-to-digital converter (ADC) at the low detector gain setting.

[0150] In this regard, 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 of the second light detector 224 corresponding only to the filtered measurement at the excitation wavelength can be considered. In various aspects, for a given light detector, the 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 fluorophore is introduced into the patient 202 as MB-102 at a dose level of about 4 μmol / kg of the patient's body weight, the expected peak detected fluorescence signal from the exogenous fluorophore during injection and renal elimination is typically expected to be about 10% of the signal magnitude 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 the high setting is still below 10% of the range of the detector ADC, the detector gain for this measurement is increased by 10 times. In another aspect, the saturation state can persist for a predefined period of time, including but not limited to a period of 30 seconds prior to adjusting the detector gain or LED power, to avoid reacting to stray signal spikes.

[0151] In another aspect, if the detected light signal from one of the light detectors 222 / 224 exceeds a threshold percentage of the maximum ADC range, the light detector control unit 232 can adjust the detector gain to a lower gain level to avoid signal saturation. Although 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 about 40% or higher, while the onset of mild detector nonlinearity occurs when the threshold percentage exceeds about 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 light detectors 222 / 224 exceeds about 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 approaching saturation, it is adjusted to low. If the current detector gain setting is low and the corresponding detected light signal remains above the threshold percentage of the maximum ADC range, the LED output power setting of the corresponding LED light source can be reduced by 10 times.

[0152] In one aspect, the optical detector control unit 232 may receive one or more feedback measurements for modulating a plurality of detector signals to compensate for changes in the performance of the optical detector due to changes in temperature and / or light source output. Non-limiting examples of feedback measurements used by the optical detector control unit 232 include: the optical output of the light sources 218 / 220 measured by the first monitor photodiode 904 and the second monitor photodiode 906, respectively, within the source well 902 (see Figure 11 ), the temperature of the optical detectors 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 related to monitoring the performance of the optical detectors 222 / 224.

[0153] In various aspects, the optical detectors 222 / 224 may be silicon photomultiplier (SPM) detectors, which may include low-noise internal amplification and may operate at lower light levels relative to other optical sensor devices such as PIN photodiodes. The detector signals generated by the SPM detectors 222 / 224 may be amplified using transimpedance amplifiers 1120 / 1118, respectively (see Figure 11 ), to convert the current generated by each SPM optical detector 222 / 224 into a measurable detector voltage. The transimpedance amplifier 1118 on the second SPM optical detector 224 (i.e., only detecting the filtered light at the excitation wavelength) may include a switchable detector gain, which may select a low gain that is configured to detect a larger dynamic range for fluorescence measurements 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 may further select a high gain setting for the second SPM optical detector 224 to enhance the sensitivity of the second SPM optical detector 224 when detecting light at the emission wavelength generated by an exogenous fluorophore within the tissue of the patient 202 during a stage of the detection cycle, to ensure that the expected dark current from the second SPM optical detector 224 occupies less than 1 / 4 of the total ADC output range. In one aspect, the second transimpedance amplifier of the second SPM optical detector 224 may include a low detector gain that is configured to provide a transimpedance gain of approximately 4 kΩ, which corresponds to approximately twice the value of the transimpedance resistor due to differential operation, and may also include a high detector gain that is configured to provide a transimpedance gain of approximately 40 kΩ. In another aspect, the first transimpedance amplifier of the first SPM optical detector 222 may include a fixed detector gain that is configured to provide a transimpedance gain of approximately 2 kΩ.

[0154] iii) Acquisition unit

[0155] Referring again to Figure 2 , in various aspects, the controller 212 may further 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, and the additional light detector 226 and the additional temperature sensor 228, and process the plurality of signals to generate one or more raw signals, which include but are not limited to, a raw fluorescence signal of 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 the two light detectors 222 / 224 during illumination at the emission wavelength.

[0156] The plurality of signals received from the various sensors and devices described above are generally 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 the circuit 1100, showing the arrangement of various electrical devices and components of the sensor head 204. In one aspect, an analog signal encoding the intensity of the light detected by the first light detector 222 and the second light detector 224 may be received by the first ADC 1102.

[0157] In various aspects, at least one 24-bit Σ-Δ ADC may be used to digitize the analog signals generated by the light detectors 222 / 224 and the various monitor sensors. Referring again to Figure 11, in one aspect, a high-speed 24-bit Σ-Δ ADC 1102 can be used to digitize an analog signal encoding measurements from a time-sensitive sensor. In this regard, time-sensitive sensors include sensors associated with the generation and detection of optical pulses characterized by potentially rapidly changing signals. Non-limiting examples of the time-sensitive sensors of system 200 include: first and second photodetectors 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 an analog signal encoding measurements from a less time-sensitive sensor. In this other aspect, less time-sensitive sensors include sensors associated with monitoring system conditions characterized by typically slowly changing signals, which system conditions include but are not limited to the temperature of various system components and / or regions. Non-limiting examples of the less time-sensitive sensors of system 200 include: first and second thermistors 1106 / 1108 configured to monitor the temperature of the light sensors 222 / 224 and the light sources 218 / 220, respectively, and a third temperature sensor 1128 configured to monitor the temperature of the housing 600 of the sensor head 204.

[0158] In various aspects, the acquisition unit 234 can also be configured to be capable of synchronously detecting light via the detectors 222 / 224. Without being limited to any particular theory, the synchronous detection method is believed to suppress noise from detector signals associated with the detection of light generated by the light sources 118 / 120 and fluorescence generated by exogenous fluorophores within the tissue of the patient 202 by differentiating the detector signals from noise associated with detecting ambient light or other sources of interference.

[0159] Figure 12 is a schematic diagram of the synchronous detection method in one aspect. Referring to Figure 11 and Figure 12 , the waveform generator / FPA 1122 can generate a digital square wave 1202 received by the DAC 1124, and the resulting analog-converted square wave is received by the LED current source 1126. The resulting current, characterized by a waveform also proportional to the analog-converted square wave, generated by the LED current source 1126 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 together with the fluorescence generated by endogenous fluorophores and digitized by the high-speed ADC 1102.

[0160] Referring again to Figure 11 and Figure 12 , the digital square wave 1202 generated by the waveform generator / FPA 1122 can also be received by the DAC1110 (see Figure 11)Converted to an in-phase reference sine wave 1210 and a 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 signed multiplied at a first multiplier 1214 to generate a plurality of in-phase modulation signals. Additionally, the digitized detector signal and the quadrature reference cosine wave 1212 can be sampled and signed multiplied at a second multiplier 1216 to generate a plurality of quadrature (out-of-phase) modulation signals. In this aspect, the acquisition unit 234 can delay the samples from the reference wave 1210 / 1214 by an amount equal to the relative delay between the DAC 1124 that generates the reference wave 1210 / 1214 and the ADC 1102 of the digitized detector signal to synchronize the reference wave 1210 / 1214 with the detector data being acquired.

[0161] Referring again to Figure 12 , the in-phase modulation signals can be summed in a first accumulator 1218 to generate an in-phase intensity signal 1224. Similarly, the quadrature modulation signals can be summed in a third accumulator 1222 to generate a quadrature intensity signal 1228. The original digitized detector signal can also be summed in a second accumulator 1220 to generate an average intensity signal 1226. Additionally, the in-phase intensity signal 1224 and the quadrature intensity signal 1228 can be square root summed to generate an amplitude signal 1230.

[0162] 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 that control the synchronous detection operate in 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 on the order of about a fraction in 10 6 of.

[0163] In one aspect, the digital square wave 1202 that modulates the LED light sources 218 / 220 and implements the synchronous detection method described above herein is generated at a frequency of about 1 kHz. Without being limited to any particular theory, the square wave is selected as the modulation waveform compared to a pure sine wave that is the same peak power level as the modulation waveform to be able to enhance the signal-to-noise ratio (SNR).

[0164] In another aspect, the acquisition unit 234 may also be configured to be able to demodulate the in-phase intensity signal 1224, the average intensity signal 1226, and the quadrature intensity signal 1228. In one aspect, the acquisition unit 234 may pick out each component at the fundamental wave, characterized in that the amplitude is (4 / π) times the amplitude of the square wave 1202 used to modulate the intensity signals 1224 / 1226 / 1228. In various aspects, in order to suppress the 50 / 60 Hz electrical noise generated by the AC power supply and the corresponding 100 / 120 Hz optical noise generated by the ambient light sources powered by these power supplies, the integration period of the accumulators 1218 / 1220 / 1222 may be selected as a multiple of 100 ms. In these different aspects, the selected integration period ensures that the integration of the accumulators 1218 / 1220 / 1222 occurs over an integer number of cycles of the 50, 60, 100, and 120 Hz signals.

[0165] iv) Processing unit

[0166] Referring again to Figure 2 , in various aspects, the controller 212 may also include a processing unit 236, which is configured to apply corrections to the demodulated detector signals and transform selected portions of the corrected detector signals into measurements of renal function. Figure 13 is a block diagram showing the subunits of the processing unit 236 in one aspect. Referring to Figure 13, the processing unit 236 may include a preprocessing subunit 1302 configured to determine and correct the detector signals to remove signal artifacts associated with various confounding effects, including but not limited to physiologically induced signal variations, power supply variations to the light sources 218 / 220, non - linearities in the detector response, ambient temperature variations, and tissue heterogeneity. The processing unit 236 may further include a baseline subtraction subunit 1304 configured to remove a portion of the detector signal attributable to external 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 diffuse reflection correction subunit 1306 configured to apply a diffuse reflection correction method to remove the effects of diffuse reflection of light within the tissue of the patient 202. The processing unit 236 may further include a post - balance selection subunit 1308 configured to select a portion of the detector data associated with a post - reagent administration period for subsequent analysis to determine the renal function of the patient. The processing unit 236 may further include an RDTC calculation subunit 1310 configured to transform the detector signals obtained during the post - reagent administration period to generate a renal decay time constant indicative of the renal function of the patient. The processing unit 236 may further include a fault detection subunit 1312 configured to monitor the magnitude of the detector signals to detect any faults in the system.

[0167] a) Preprocessing subunit

[0168] In one aspect, various modules of the preprocessing subunit 1302 are used to preprocess the raw signals corresponding to the light intensities detected by the light detectors 222 / 224, where the light intensities detected by the light detectors 222 / 224 correspond to the irradiations of the first light source 218 and the second light source 220 at the excitation wavelength and the emission wavelength, respectively, to remove the effects of multiple confounding factors from the raw signals, thereby generating signals that more accurately reflect the underlying signal of specific interest.

[0169] As multiple non - limiting examples, the intensity of the light generated by the light source may vary due to one or more of a plurality of factors, including but not limited to: fluctuations in the current supplied to the light source and variations in the ambient temperature of the light source. Light characterized by two or more wavelengths emitted from the same source aperture of the sensor head may not share the same path to the same detector. The detector may have temperature - dependent sensitivity and gain. Additionally, the optical filter associated with the second light detector 224 may have temperature - dependent transmission characteristics.

[0170] In one aspect, the preprocessing subunit 1302 is configured to process the raw signals corresponding to the light intensities detected by the first and second photodetectors 222 / 224 to remove one or more of the measurement errors associated with the devices and components of the system 200 and patient-specific factors, where the patient-specific factors include, but are not limited to, the above-mentioned multiple factors. Figure 22A is a block diagram of a module showing the preprocessing subunit 1302 in one aspect. Figure 22B is a block diagram of a module showing the preprocessing subunit 1302a in a second aspect.

[0171] In one aspect, as Figure 22A shown, the preprocessing subunit 1302 1) resamples the signal using the method of the resampling module 2202 described below, 2) removes the saturated detector signal using the method of the detector output saturation detection and removal 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 the instrument light directionality using the method of the light directionality correction module 2208 described below, 5) corrects the signal for the filter throughput and temperature-dependent changes in fluorescence using the method of the filter throughput temperature correction (emission) module 2212 described below, 6) corrects the tissue non-uniformity using the method of the tissue non-uniformity correction module 2216 described below, 7) corrects the signal for the filter throughput and temperature-dependent changes in the excitation light and signal decomposition using the method of the filter throughput temperature correction (excitation) module and signal decomposition module 2214 described below, 8) corrects the light power variation using the method of the fractional photon normalization module 2218 described below.

[0172] In one aspect, as Figure 22B shown, the preprocessing subunit 1302a calculates the signal magnitude using the method of the detector temperature correction module 2206a described below, resamples the signal using the method of the resampling module 2202a described below, removes the saturated samples using the method of the detector output saturation detection and removal module 2204a described below, corrects the signal for the temperature-dependent detector gain using the method of the detector temperature correction module 2206a described below, corrects the signal for the light power variation using the method of the fractional photon normalization module 2218a described below, corrects for excitation light leakage into the measured fluorescence signal using the filter throughput temperature correction (excitation) module and signal decomposition module 2214a described below, and corrects for fluorescence leakage into the measured excitation diffuse reflection signal using the filter throughput temperature correction (emission) module 2212a described below.

[0173] - Resampling module

[0174] Referring toFigure 22A and Figure 22B ,in various aspects, the preprocessing subunit 1302 / 1302a includes a resampling module 2202 / 2202a, and the resampling module 2202 / 2202a is configured to reduce signal variations associated with the physiological processes of patient 202, which include but are not limited to heartbeat and respiration. Generally, the acquisition sequence is characterized in that the alternating illumination intervals at excitation and emission are separated by non-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 intervals between physiological events such as heartbeat or respiration, physiological noise may 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 before subsequent processing of the signals.

[0175] As a non-limiting example, the sample sequence can 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. Since physiological signal variations, such as those from the heartbeat, occur on the same time scale, the 200 ms difference between the signal acquisitions associated with the excitation and emission wavelengths becomes apparent in the signal. The preprocessing subunit 1302 can be used to reduce this physiological signal noise by first resampling the signals associated with the illuminations at the excitation and emission wavelengths to overlap before performing any additional signal processing as described below. In this non-limiting example, the signal associated with the illumination at the excitation wavelength can be shifted forward by 100 ms, and the signal associated with the illumination at the emission wavelength can be shifted backward by 100 ms, resulting in an overlap of the signals.

[0176] In various aspects, the resampling module 2202 performs the resampling as described above on the signals detected by both the first and second detectors 222 / 224. In one aspect, the resampling module 2202 is in the form of a low-pass filter.

[0177] - Detector output saturation detection and removal module

[0178] Referring again to Figure 22A and Figure 22B, in various aspects, the preprocessing subunit 1302 / 1302a includes a detector output saturation detection and removal module 2204 / 2204a, which is configured to detect and remove signal values that fall outside the detection range of the photodetector 222 / 224. In one aspect, the preprocessing subunit 1302 compares the detected signal with the maximum ADC signal. If any signal falls within the threshold range of the maximum ADC signal using the average value or peak signal value, the detector output saturation detection and removal module 2204 identifies the value and removes the value from further processing.

[0179] - Detector temperature correction module

[0180] Figure 22A and Figure 22B , in various aspects, the preprocessing subunit 1302 / 1302a includes a detector temperature correction module 2206 / 2206a, and the detector temperature correction module 2206 / 2206a is configured to be able to perform temperature correction to compensate for the thermal sensitivity of the photodetector 222 / 224. In one aspect, the intrinsic detector gain of a silicon photomultiplier (SPM) device, which is commonly used as a photodetector, is proportional to the difference between the device breakdown voltage and the bias voltage applied by the bias voltage generator 1112 (see Figure 11 ). In this regard, the breakdown voltage changes with temperature in a well-characterized manner. In one aspect, the temperature correction takes into account both this internal detector gain change and the additional temperature-related changes in the photon detection efficiency.

[0181] 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 the calculated gain G(t) to remove the temperature dependence. The scaling correction G(t) can be calculated according to Equation (2):

[0182]

[0183] In Equation (2), the monitored temperature T is obtained from a first temperature sensor 1106 (see Figure 11 ) configured to monitor the temperature of the sensors 222 / 224. The 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 breakdownIt can be 24.5V and T0 can be 21°C. On the other hand, the coefficients C used in Equation (2) can be empirically derived based on measurements obtained using a constant phantom within an ambient temperature range of approximately 18°C to approximately 26°C v and C T .

[0184] On the other hand, the temperature part of the gain correction is determined by Equations (3) to (5).

[0185]

[0186]

[0187] This gain correction can be applied to each signal amplitude measured by the first and second photodetectors 222 / 224 as follows:

[0188]

[0189] In one aspect, according to Equation (1), the amplitude of the temperature-corrected measurement from each detector and monitor photodiode is calculated based on the square root sum of the in-phase intensity signal 1224(I) and the quadrature intensity signal 1228(Q):

[0190]

[0191] The signal amplitude from the photodetectors 222 / 224 calculated using Equation (1) is normalized by the monitor photodiode amplitude for each measurement group corresponding to measurements obtained during illumination at the excitation wavelength or emission wavelength of one of the LED light sources 218 / 220. In one aspect, if one photodiode is located in the source well 902, the single photodiode amplitude from the corresponding measurement group is used for this normalization. On the other hand, if two monitor photodiodes 904 / 906 are located in the same source well 902 as two LED light sources 218 / 220 (see Figure 9 ), the average of the two monitor photodiode amplitudes from the corresponding measurement group is used for this normalization.

[0192] 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 such that the intensity signals 1224 / 1226 / 1228 are returned as part of the entire range of the high-speed ADC 1102 (i.e., the range from a minimum of 0 to a maximum of 1). The measurements of the monitor photodiodes 904 / 906 (see Figure 11 ) are similarly scaled to be part of the entire range of the low-speed ADC 1104.

[0193] In one aspect, G correction may incorporate power correction to correct for the effects of fluctuations in the LED power supply. In this regard, when the light intensity from the light sources 218 / 220 changes, 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. For each light source 218 / 220, C source1 and C source2 's calibration coefficients are calculated as the milliwatts measured by the detector for each recorded monitor photodiode signal value. C source1 and C source2 are used to determine the absolute light output into the tissue at each wavelength.

[0194] Referring again to Figure 22B , the detector temperature correction module 2206a corrects the signal amplitude for the varying intensity of the LED by normalizing the temperature-corrected detection signal using the LED output signal PD measured by the first monitor photodiode 904 and / or the second monitor photodiode 906. In this case, for each light source 218 / 220 from above, the G magnitude variable is modified as follows: correction - Light directionality correction module

[0195]

[0196] - Light directionality correction module

[0197] Referring again to Figure 22A , in this regard, the preprocessing subunit 1302 includes a light directionality correction module 2208, which is configured to be able to correct for variations in the detection signal associated with differences in the scattering and absorption of light of different wavelengths through the tissue of the patient 202 during data acquisition. In one aspect, the correction term for light directionality can be measured by obtaining data from one or more homogeneous tissue models and using a sensor configuration without an emission filter. 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 the coefficient G ex or G em of the signal, and the coefficient G ex or G em associated with the illumination of light at the excitation wavelength and the emission wavelength is obtained respectively. 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 coefficient G ex or G emThe correction of the signals acquired in a homogeneous medium makes the signals measured by the first detector 222 / 224 and the second detector 222 / 224 equivalent within 20% of each other. In other respects, the first optical detector 222 uses a coefficient G ex or G em The correction of the signals acquired in a homogeneous medium makes the signals measured by the first detector 222 / 224 and the second detector 222 / 224 equivalent to approximately 10%, approximately 5%, approximately 2%, and approximately 1%.

[0198] - Detector Nonlinear Response Correction Module

[0199] Referring again to Figure 22A , in this aspect, the preprocessing subunit 1302 includes a detector nonlinear response correction module 2210, which is configured to be able to correct the variations in the detection signals associated with the nonlinear response of the detector. In this regard, a calibration curve based on average data can be used to scale the amplitude data obtained by the detectors 222 / 224.

[0200] - Filter Throughput Temperature Correction (Transmission) Module

[0201] Referring again to Figure 22A , in this aspect, the preprocessing subunit 1302 includes a filter throughput temperature correction (transmission) module 2212, which is configured to be able to correct the variations in the detection signals associated with the temperature-dependent optical characteristics of the optical filter 244 associated with the second optical detector 224 during the emission wavelength illumination. In this regard, the signal Det2 detected by the second optical detector 224 can be corrected according to Equation (8):

[0202]

[0203] In various aspects, when the ambient temperature is cycled within a range including the operating temperature range or a sufficiently large subset of the range to fully determine the temperature dependence of the emission filter, the signal Det2 measured by the second optical detector 224 can be monitored. These data are obtained from a homogeneous, non-fluorescent phantom using the optical filter 244 mounted on the second optical detector 224. In addition, simultaneous measurements are monitored from the first optical detector 222, and the ratio of the measured values Det2 / Det1 is determined. The nominal filter coefficient C emF,nom is calculated as the nominal ratio of Det2 / Det1 obtained at the nominal operating temperature T nom . In this regard, the coefficient C emF,slopeT is obtained from the slope of Det2 / Det1 obtained during the emission wavelength illumination of the homogeneous non-fluorescent phantom within the ambient temperature range.

[0204] - Tissue Heterogeneity Correction Module

[0205] Referring again to Figure 22A , in this regard, the preprocessing subunit 1302 includes a tissue heterogeneity correction module 2216, which is configured to be able to correct the variation of the detection signal associated with the heterogeneity of the tissue intervening between the first region 206 irradiated by the light sources 218 / 220 and the second and third regions 208 / 210 in which the photodetectors 222 / 224 are located. In this regard, according to equation (9), the signal Det1 corrected to light directionality by the light directionality correction module 2208 and the signal Det2 corrected to filter effect by the filter throughput temperature correction (emission) module 2212 are used to calculate C hetero , that is, the coefficient for correcting tissue heterogeneity:

[0206] C hetero = Det2 / Det1 Equation (9)

[0207] - Filter throughput temperature correction (excitation) and signal decomposition module

[0208] Referring again to Figure 22A , in this regard, the preprocessing subunit 1302 includes a filter throughput temperature correction (excitation) module and a signal decomposition module 2214. The filter throughput temperature correction (excitation) module and the signal decomposition module 2214 are configured to be able to correct the variation of the detection signal associated with the temperature-dependent optical characteristics of the optical filter 244 associated with the second photodetector 224 during the illumination at the excitation wavelength. In this regard, since the emission filter is configured to block the light at the excitation wavelength, the filter throughput temperature correction (excitation) module and the signal decomposition module 2214 perform the correction of the variation of the excitation light leakage amount due to the temperature-related change of the optical properties of the optical filter 244. In addition, the filter throughput temperature correction (excitation) module and the signal decomposition module 2214 can correct the signal measured by the first photodetector 222 during the illumination at the excitation wavelength due to the presence of fluorescence caused by the excitation wavelength illumination superimposed on a part of the signal associated with the illumination at the excitation wavelength.

[0209] In this regard, as shown in equation (10), the influence of the temperature-dependent change on the leakage of the excitation wavelength through the optical filter 244 is calculated:

[0210] C exLT =C exLT,nom +C exLT,slopeT (T - T nom ) Equation (10)

[0211] In this regard, during the illumination at the excitation wavelength at the nominal operating temperature T nomCalculate C from the ratio of the signals Det1 and Det2 measured from a uniform non-fluorescent phantom exLT,nom C exLT,slopeT is calculated as the slope of the signal Det2 measured from a uniform non-fluorescent phantom within the operating temperature range T during excitation wavelength illumination.

[0212] In this regard, the filter throughput temperature correction (excitation) module and the signal decomposition module 2214 also perform signal extraction to isolate the portion of the detection signal associated with the diffuse reflection of the excitation wavelength illumination and fluorescence. DR ex2 , i.e., the amount of excitation light incident on the second photodetector 224 in the absence of the optical filter 244, is not measurable due to the presence of the optical filter 244. Additionally, the signal Det1 measured by the first photodetector 222 is a composite signal from both the diffuse reflection and fluorescence Flr1 of the excitation wavelength illumination DR ex1 . C is obtained using the tissue heterogeneity correction module 2216 as described above Hetero . The baseline signal is extracted using the following system of equations:

[0213] Det2 = C exLT DR ex2 + Flr2 Equation (11)

[0214] Det1 = DR ex1 + Flr1 Equation (12)

[0215] Flr2 = C Hetero Flr1 Equation (13)

[0216] DR ex2 = C Hetero DR ex1 Equation (14)

[0217] In this regard, Flr2 is determined by solving the above system of equations using only the measurable signals Det1 and Det2, as follows:

[0218] Det2 = C exLT C Hetero DR ex1 + Flr2 Equation (15)

[0219] Det2 = C exLT C Hetero (Det1 - Flr1)+ Flr2 Equation (16)

[0220] Det2 = C exLT C Hetero Det1 - C exLT C Hetero Flr1 + Flr2 Equation (17)

[0221] Det2-C exLT C Hetero Det1 = Flr2(1 - C exLT ) Equation (18)

[0222]

[0223] In this regard, once Flr2 is obtained as described above, other signals Flr1, DR ex1 and DR ex2 .

[0224] - Partial photon normalization module

[0225] Referring again to Figure 22A , in this regard, the preprocessing subunit 1302 includes a partial photon normalization module 2218, which is configured to convert the detector signal into partial photon units after the preprocessing 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 inverting the scaling associated with the ADC and the transimpedance amplifier, and the transimpedance amplifier is used to acquire the detected signal to obtain the 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 watts. Then, the detector signal in watts is proportional to the source power in watts, which is measured by an additional photodetector 226 for monitoring the output of the light source 218 / 220, to obtain the number of detected partial photons.

[0226] - Optical power correction module

[0227] Referring again to Figure 22A and Figure 22B, in this regard, the preprocessing subunit 1302 / 1302a includes a partial photon normalization module 2218 / 2218a, which is configured to convert the detector signal into partial photon units after the preprocessing 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 a photocurrent by inverting the scaling associated with the ADC and the transimpedance amplifier, which is used to acquire the detected signal to obtain the 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 watts. Then, the detector signal in watts is proportional to the source power in watts, which is measured by an additional photodetector 226 for monitoring the output of the light source 218 / 220, to obtain the number of detected partial photons.

[0228] - Excitation light leakage subtraction module

[0229] Referring again to Figure 22B , in this regard, the preprocessing subunit 1302a includes a partial photon normalization module 2222, which is configured to perform excitation leakage subtraction on the Flr meas signal. To obtain the fluorescence signal generated only by fluorescence photons (Flr photons ), excitation leakage subtraction is performed. To remove the contribution of the excitation light, the excitation leakage is considered as part of the diffuse reflection excitation signal, where a general calibration factor, C ExLT , determines the part of the signal to be subtracted from Flr meas as follows:

[0230]

[0231] where C ExLT is a calibration factor obtained by calculating the ratio of the excitation light detected by two detectors on a non-fluorescent optical model as described below:

[0232]

[0233] Then this signal is subtracted from Flr meas to provide a fluorescence signal generated only by fluorescence photons as follows:

[0234] Flr photons = Flr meas - ExLT

[0235] - Fluorescence light leakage subtraction module

[0236] Referring again toFigure 22B , in this aspect, the preprocessing subunit 1302a includes a fluorescence light leakage subtraction module 2224a configured to perform fluorescence light leakage subtraction on the Flr meas signal. To obtain diffuse reflection, which is defined herein as the excitation signal caused only by excitation photons (DRex photons ), fluorescence light leakage subtraction is performed. To remove fluorescence light leakage, a calibration factor, C , is determined based on the relationship between the amount of fluorescence light leakage observed on the database of human subject data and the tissue heterogeneity measured between the diffuse reflection and emission signals FlrLT . This relationship is a linear relationship as shown below:

[0237]

[0238] where in one aspect p1 and p2 are approximately 0.61 and 0.01, respectively, as determined by the above relationship. In another aspect, without being limited to that defined by the above relationship, p1 and p2 can assume any other values.

[0239] Then, the DRex photons signal is calculated by subtracting this portion of the measured fluorescence from the diffuse reflection excitation signal, as follows:

[0240]

[0241] b) Baseline subtraction subunit

[0242] Referring again to 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 photodetector measurements to correct for the effects of autofluorescence and light leakage. The baseline period used herein refers to the initial time period of the measurements obtained before the injection of the exogenous fluorescent agent. During the baseline period, it can be assumed that the fluorescence signal measured by the system 200 is associated with the tissue autofluorescence and / or excitation light from the LED light sources 218 / 220 that leaks through the absorption filter 244 of the second photodetector 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 measurements associated only with the fluorescence generated by the exogenous fluorescent agent within the patient's tissue.

[0243] In another aspect, the correction for excitation light leakage and autofluorescence can be implemented separately. In this other aspect, the subtraction of the excitation light leakage effect can be performed before the diffuse reflection correction described below herein, and the subtraction of the autofluorescence effect can be performed after the diffuse reflection correction.

[0244] c) Diffuse reflection correction subunit

[0245] Referring again to Figure 13 , the processing unit 236 further includes a diffuse reflection correction sub-unit 1306. In one aspect, the diffuse reflection correction sub-unit 1306 can correct the measured fluorescence data to remove the effects of changes in the optical properties (absorption and scattering) of the tissue of patient 202 during the monitoring of the renal extraction of the exogenous fluorophore within the patient tissue. As described above, the optical properties of the tissue can be altered due to any one or more factors, including but not limited to: vasodilation, vasoconstriction, blood oxygen saturation, hydration, edema, and any other suitable factors within the region of interest being systematically monitored that are associated with changes in the concentration of endogenous chromophores such as hemoglobin and melanin.

[0246] 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. In response to illumination by the excitation wavelength photons, these emission wavelength photons are emitted by the exogenous fluorophore introduced into the patient tissue. The emission wavelength photons travel from the fluorescence source (i.e., the exogenous fluorophore) 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 can also include autofluorescence emitted by endogenous fluorophores (such as keratin and collagen) within the patient tissue, as well as leakage of the excitation wavelength light through the optical filter 244 of the second light detector 224. The excitation wavelength photons that induce the fluorescence of the exogenous fluorophore 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 during the time interval in which the detector data for determining renal function are being acquired (i.e., from a few hours to about 24 hours or longer), the accuracy of the fluorescence measurements can be affected, as discussed above.

[0247] During each measurement cycle in one aspect, the system 200 can direct light into the 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 the light emitted from the second region of the patient's skin using the first (unfiltered) light detector 222, and detect a portion of the light emitted from the third region 210 of the patient's skin using the second (filtered) light detector 224. The light intensity detected by each combination of excitation and emission wavelength illumination of the first region 206 and detection by the unfiltered / filtered light detectors 222 / 224 contains information not only about the concentration of the exogenous fluorophore in the patient tissue, but also about the optical properties of the patient's skin.

[0248] Table 2: Photodetector measurements corrected for temperature and power fluctuations

[0249]

[0250]

[0251] The main measurement of fluorescence is Flr meas , the intensity of fluorescence measured at the filtered detector.

[0252] Diffuse reflectance measures Flr meas represents the propagation of photons to the unfiltered device and consists mainly of excitation photons.

[0253] DR em and DR em,filtered represent the propagation of only emitted photons.

[0254] Referring to Table 2, in each aspect, the light intensity measured by the second (filtered) light detector 224 during irradiation by the excitation wavelength light, before any correction of the tissue optical properties is captured, is the raw intensity of the light emitted by the exogenous fluorophore (Flr meas ). After the baseline subtraction correction as previously described herein, assuming that the emission wavelength light contained in Flr meas originates mainly from the exogenous fluorophore, with only a small contribution due to the autofluorescence of endogenous chromophores, and is thus 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.

[0255] However, if the optical properties of the patient's skin change during data acquisition, the patient's skin may exhibit more or less autofluorescence at the emission wavelength, thus introducing uncertainty into the accuracy of the previously performed background subtraction correction. In addition, the changing skin optical properties can further alter the intensity of the light reaching the exogenous fluorophore at the excitation wavelength, in response to irradiation by the excitation wavelength light, thereby changing the amount of energy absorbed by the exogenous fluorophore and the intensity of the induced fluorescence emitted from the exogenous fluorescence. In each aspect, the remaining three light measurements can monitor the optical properties of the patient's skin and provide data that can be used to adjust for any changes in the optical properties of the patient's skin.

[0256] Referring again to Table 2, before applying the diffuse reflectance correction, the signals DRex meas and Flr meas , which have been corrected for changes in temperature and optical output, are further processed into signals that are only attributable to photons of the desired wavelength. Since the number of photons due to diffuse reflectance, excitation, or fluorescence on any detector depends on the light directionality and the gain of the detector at the detection wavelength, as follows:

[0257] DRex meas= A1 * DRex photons + B1 * Flr photons

[0258] Flr meas = A2 * DRex photons + B2 * Flr photons

[0259] where the coefficients A1, A2, B1, and B2 consist of a directionality and a gain factor. For example,

[0260] A1 = d450 SPM1 * G SPM1@450

[0261] The isolation of the signal generated from fluorescence emission and diffuse reflection, at the excitation wavelength photons, is performed as follows:

[0262]

[0263] Since the renal function monitor measures a rate independent of amplitude, no photon signal is required (e.g., the previous constant term, as shown below.

[0264]

[0265] Thus, the terms (or C ExLT ) and (or C FlrLT ) can be experimentally determined to separately isolate Flr photons and DRex photons .

[0266] The following table shows the names of the signals used to represent each of the four measurement signals in the diffuse reflection correction expansion. Note that either of the described preprocessing paths can be followed to obtain the signals that can be used for the expansion correction.

[0267] Table 3: Photodetector measurements for obtaining fluorescence measurements corrected for variable tissue optical properties

[0268]

[0269] where any one of the excitation wavelength signals can be used as an alternative method to obtain the diffuse reflection correction through any one of the described preprocessing methods.

[0270] Referring again to Table 2, the light intensity measured by the first (unfiltered reference) photodetector 222 during illumination by the excitation wavelength light captures the measured value of the diffuse reflection of the excitation wavelength light propagating through the patient's skin Although the first light detector 222 is configured to detect both the excitation wavelength light and the emission wavelength light, due to the relatively low efficiency of light generation via fluorescence, the intensity of the excitation wavelength light is many orders of magnitude higher than the intensity of the emission wavelength light. In various aspects, it is assumed that the proportion of light at the emission wavelength in is negligible. In other aspects, the proportion of the emission wavelength light in is estimated and subtracted. Without being limited to any particular theory, since the intensity of the excitation wavelength light guided to the patient's skin is assumed to be relatively constant and the loss due to absorption by the exogenous fluorophore is negligible, and power correction is performed as previously described herein, is used as a reference measurement to evaluate changes in the optical properties of the patient's skin relative to the excitation wavelength light.

[0271] The light intensity measured by the first (unfiltered reference) light detector 222 during illumination by the emission wavelength light captures a measurement of the diffuse reflection of the emission wavelength light propagating through the patient's skin (DR em ). Without being limited to any particular theory, since no emission wavelength light is induced from the exogenous fluorophore due to the absence of excitation wavelength illumination during this stage of the data acquisition cycle, and since the intensity of the emission wavelength light guided to the patient's skin is relatively constant and power correction is performed as previously described herein, DR em is used as a reference measurement to evaluate changes in the optical properties of the patient's skin relative to the emission wavelength light.

[0272] The light intensity measured by the second (filtered) light detector 224 during illumination by the emission wavelength light captures a second measurement of the diffuse reflection of the emission wavelength light propagating through the patient's skin (DR em,filtered ). In one aspect, the same assumptions are made for DR em,filtered as for the above DR em . Additionally, DR em,filtered provides a means to evaluate the heterogeneity of tissue optical properties. Since DR em,filtered 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 ), the intensity of the light measured in DR em,filtered has propagated along an optical path through the patient's skin that is different from the optical path traveled by the light measured in DR em . Without being limited to any particular theory, since the distances of the first detector aperture 1004 and the second light aperture 2006 (through which light is transmitted to the first and second light detectors 222 / 224) are designed to be equidistant from the light transmission aperture 1002 respectively (see Figure 10 ), it is assumed that DRem,filtered and DR em Any difference between them is caused by the heterogeneity of the optical properties of the skin through two different optical paths.

[0273] In one aspect, intrinsic fluorescence (IF), as defined herein as the measured fluorescence at the emission wavelength that is solely attributable to the emission of exogenous fluorophores, can be calculated according to Equation (20):

[0274]

[0275] Factors IF, Flr, DR ex 、DR ex and DR em,filtered are defined as above herein. As shown in Equation (20), the diffuse reflectance corrected measurement signal DR ex 、DR em and DR em,filtered Each of the factors is raised to the power k ex 、k em and k em,filtered . In one aspect, prior to applying the diffuse reflectance correction of Equation (20), each measurement in Table 2 is subjected to the power / temperature correction and background subtraction correction as described above (see Figure 22A and Figure 22B ).

[0276] In various aspects, k ex 、k em and k em,filtered can be determined empirically. Non-limiting examples of suitable empirical methods for determining the suitable values of k ex 、k em and k em,filtered include the global error mapping method and the linear regression method, both of which are described in detail below.

[0277] In one aspect, once each power (k ex 、k em 、k em,filtered) For the values, the same set of exponents can be reused for subsequent autofluorescence measurements. Non-limiting examples of applications of the systems and methods described herein where a selected set of exponents can be reused include: repeated measurements on the same patient using the same sensor head 204; repeated measurements on different patients of the same species using the same sensor head; repeated measurements on patients of the same species using different sensor heads of the same design; repeated measurements on patients of the same species using different sensor heads of different designs; repeated measurements on patients of different species using different sensor heads of different designs; and any other suitable applications of the systems and methods described herein. In another aspect, the exponents can be updated by reusing the systems and methods described herein. In this other aspect, a new set of exponents can be determined for each use of the system and method, and the stored set of exponents can be evaluated periodically or continuously to evaluate whether an updated selection of exponents is indicated. As a non-limiting example, if analysis of multiple sets of exponents determines that the exponents did not vary outside a threshold range in a previous use of the system, the system can be used to make measurements using a previous set of exponents, the average / median of all previous sets of exponents, or any other estimate of a suitable exponent based on the previous exponent values. In this non-limiting example, if analysis of multiple sets of previous exponents determines that the exponents vary outside the threshold range, re-selection of exponents using one of the methods described below can be indicated.

[0278] Global error mapping method

[0279] In one aspect, the value of the power used in equation (20) above is determined empirically using the global error surface method.

[0280] In Figure 14A A flowchart is shown that illustrates the various steps of the global error surface method 1400. In this aspect, the method includes, at step 1402, selecting a range of values for each power (k ex , DR em , DR em,filtered ) for each diffuse reflectance signal (DR ex , k em , k em,filtered ) selected by the user. In various aspects, the range of values for each power can be affected by any one or more of a variety of factors, including but not limited to: the design of the system 200, including the design of the sensor head 204; the nature of the selected exogenous fluorophore, such as the excitation wavelength / emission wavelength, absorption efficiency, emission efficiency, and initial dose concentration in the patient tissue; the species of the patient 202 and the corresponding concentrations of endogenous chromophores; the location of the sensor head 204 on the patient 202; and any other relevant factors.

[0281] In one aspect, the method can include, for each coefficient (kex , k em , k em,filtered ) Select a wide range and conduct an extensive search. The error surface from this extensive search can be analyzed to locate wells in the error surface and the associated ranges for each coefficient. In this regard, the method includes adjusting the range of each coefficient to include the regions from the wide search, observing the wells in the error surface within this wide search, and repeating the analysis. This method can be iterated until an appropriate fine resolution that can accurately obtain the minimum error is obtained. In a non-limiting example, for a human patient, the selected range of the latent factor can be k ex in the range of [0, 2], k em in the range of [0, 4], and k em,filtered in the range of [-4, 0].

[0282] Referring again to Figure 14A , at 1404, for each power k ex , k em , k em,filtered select a step size for the range of values selected at 1402. In one aspect, the step size for each factor can be selected based on any one or more of at least several factors, which include but are not limited to: the expected sensitivity of the change in the IF value calculated by equation (20) for each factor; the total number of appropriate power combinations for calculating the IF of the factor being considered (including available computing resources, acceptable number of data processing times, or any other relevant factor); and any other suitable criteria for the step size.

[0283] In various aspects, for all powers k ex , k em , k em,filtered , the step size can be the same value. As a non-limiting example, the step size for all powers can be 0.5. In various other aspects, for all values of a single power k ex , k em , k em,filtered , the step size can be constant, but the step size selected for each power can be different between different powers. As a non-limiting example, the selected step size for k ex can be 0.01, and the selected step sizes for k em and k em,filtered can be 0.6. In various additional aspects, the step size within one or more powers can vary within the range of values of each power. As a non-limiting example, the selected step size for k ex can be non-linearly distributed with respect to the mean. In this non-limiting example, k exThe vector of potential values can be [0 0.5 0.75 0.9 1 1.1 1.25 1.5 2]. In these various additional aspects, the step size can be reduced within a sub-range of the power values for which the IF prediction calculated by Equation (20) is more sensitive to small changes at that power. Non-limiting examples of appropriately varying the step size within a single power value range include: different step sizes selected by the user, random step sizes, linear increase and / or decrease of the step size, non-linear distributions of different step sizes, such as logarithmic distribution, exponential distribution, or any other appropriate non-linear distribution of the step size.

[0284] Referring again to Figure 14A shown, the exponential range selected at 1402, and the step size selected at step 1404 can be used to form k at 1406 ex 、k em 、k em,filtered vector of potential values. As a non-limiting example, assume k ex is [0, 2], k em is [0, 4] and k em,filtered is [-4, 0] for the selected range of potential exponents, and assuming a constant step size of 0.5 for all powers, the vector created at 1406 is:

[0285] k ex = [0.0 0.5 1.0 1.5 2.0] (5 values)

[0286] k em = [0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.0] (9 values)

[0287] k em,filtered = [-4.0 -3.5 -3.0 -2.5 -2.0 -1.5 -1.0 -0.5 -0.0] (9 values)

[0288] Referring again to Figure 14A , for each combination of exponents in all vectors formed at 1406, at 1408 Equation (20) is used to calculate the IF based on the measured values Flr, DR ex 、DR em 、DR em,filtered and DR. For each combination of exponents, multiple IF values are calculated at 1408, where each IF value corresponds to one of the data acquisition cycles (i.e., as Figure 5 shown, a single sequence of emission wavelength illumination followed by excitation wavelength illumination). As a non-limiting example, using the vectors of potential exponents listed above, a total of 405 (5 * 9 * 9) IF signals will be calculated.

[0289] In one aspect, combinations of multiple potential indices can be evaluated to select a combination of indices from among the multiple to assign diffuse reflectance correction for subsequent calculation using Equation (20). Referring again to Figure 14A , an error estimate of the corrected Flr signal data calculated at 1410 (i.e., the IF signal data calculated using Equation (20)) can be calculated. Any estimate of the error can be calculated at 1410, including but not limited to curve fitting with respect to the IF signal data, quantities related to the residuals 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 an exogenous fluorophore (such as MB-102) from the kidney is expected to be a constant exponential decay, which is characterized by the renal decay time constant RDTC.

[0290] In one aspect, a subset of the Flr signal corresponding to the post-reagent administration period 1508 / 1510 can be selected to estimate the error of each combination of indices used to calculate the IF signal using Equation (20) relative to a reference curve, including but not limited to curves obtained using plasma measurements. As a non-limiting example, if an exogenous fluorophore is introduced into a patient's tissue by intravenous injection, the post-reagent administration period includes the period after injection during which the exogenous fluorophore has undergone sufficient diffusion from the blood into the extracellular fluid space of the entire patient such that the decay of the fluorescence represents the clearance of the reagent by the kidneys. In various aspects, the post-reagent administration period 1508 / 1510 of the Flr measurement can be selected by any suitable method without limitation. Non-limiting examples of suitable methods for identifying the post-reagent administration period include: selection via user inspection and automated selection methods, such as the equilibrium detection method enabled by the equilibrium selector unit 1308, as described in detail below.

[0291] Figure 15A is a graph of fluorescence measurements obtained from a patient within a period of about 10 hours after injection of an exogenous fluorophore (MB-102) approximately 3 hours before injection. Referring to Figure 15A , the pre-injection / baseline period 1502 is characterized by a relatively low and stable fluorescence level, presumably due to the absence of endogenous fluorophores in the patient's blood. After injection of the exogenous fluorophore 1503, the fluorescence measurement shows a sharp increase 1504 to a peak concentration 1506, followed by a relatively smooth exponential decay 1508 back to the background fluorescence level as the kidneys eliminate the exogenous fluorophore from the patient's blood. Without being limited to any particular theory, once the fluorescence decay is well described by a linear fit (or a line on a semi-logarithmic curve), it is believed that the injected exogenous fluorophore is likely to be in equilibrium on the extracellular space. Figure 16 is Figure 15A an enlargement of the graph ofFigure 16 Shows a comparison of the measured fluorescence data with the linear curve fit 1604 and the logarithm of the IF signal within a portion of the post-equilibrium period 1510, showing a tight fit of the mono-exponential curve fit to the IF signal data.

[0292] In one aspect, the logarithm of the calculated IF signal value can be line-fitted, and for each of a plurality of combinations of exponents, the error of the curve fit relative to the respective IF values can be compared to the IF signal calculated using Equation (20) to calculate the error at 1410. Any statistical summary parameter suitable for quantifying the error of the mono-exponential curve fit and the corresponding IF signal values can be used without limitation, including but not limited to: root mean square (RMS) error, mean absolute deviation, mean signed deviation, root mean square deviation, and any other suitable statistical summary parameter. In one aspect, the error calculated at 1410 can be the normalized RMS error of the linear fit of the logarithm(IF) signal. In this regard, the normalized RMS error calculated at 1410 is a single numerical quantity to facilitate subsequent selection of a single combination of exponents from the plurality of combinations identified at 1406.

[0293] Referring again to Figure 14A , method 1400 includes, at 1412, selecting a single combination of exponents from the plurality of combinations for which the IF is calculated at 1408. Without being limited to any particular theory, it is assumed that the combination of exponents associated with the calculated IF signal that minimizes the error calculated at 1410 is most suitable for correcting the measured Flr signal to eliminate the effect of changes in the optical properties of the patient's skin during data acquisition within the post-reagent application period 1508 / 1510. In various aspects, any known method of identifying a combination of exponents can be used without limitation, including but not limited to, selecting a single combination of exponents from a mapping of all error values corresponding to all combinations of exponents.

[0294] In various aspects, the plurality of error values corresponding to the plurality of combinations of exponents can be converted into an error map that includes a three-dimensional volume, where each of the three dimensions corresponds respectively to the powers used in Equation (20): k ex 、k em and k em,filtered . In these different aspects, each error value corresponding to one of the combinations of exponents is mapped to coordinates (k ex1 , k em1 and k em,filtered1 ) within the three-dimensional volume, where k ex1 、k em1 and k em,filtered1is a numerical value of a combination of exponents. In various aspects, each error value can be mapped to a three-dimensional volume in any known form, which includes but is not limited to: numbers, colors, grayscale values, and any other suitable form.

[0295] In one aspect, the three-dimensional mapping of the above error values can be transformed into a plurality of error surfaces, which correspond to the planar mapping of the error values associated with a single value of one of the powers k ex 、k em and k em,filtered , and the entire numerical range of the remaining two exponents is used as the horizontal and vertical axes of the error mapping.

[0296] Figure 17 is the error mapping of the normalized RMS error of the single-exponent curve fitting of the calculated IF signal, which is mapped to the entire range of k em with a constant value of k em,filtered (horizontal axis) and k ex (vertical axis), where the normalized RMS error is represented as color on the error mapping. In one aspect, the normalized RMS values calculated for each coefficient can be normalized according to Equation (21):

[0297]

[0298] where IF agent is the calculated IF signal, and fit(IF agent ) is the corresponding value of the single-coefficient curve fitting equation. In one aspect, by analyzing a single measurement group as described above, the global error mapping method determines the power for correction of changes in skin optical properties. In another aspect, the global error mapping method can analyze and combine multiple measurement data sets of multiple individuals obtained using the same system and / or sensor head. In yet another aspect, the global error mapping method can analyze multiple measurement data sets of multiple individuals obtained from using multiple systems and sensor heads. In one aspect, the power used in the correction according to Equation (20) can be determined for each measurement of each individual. In other aspects, at least several different measurement data sets can be used to obtain the power to be used, and the power thus obtained can be stored for subsequent use with measurement data sets obtained from new individuals and / or obtained using different systems and / or sensors. In various aspects, the projections across each error surface (k em,filtered -k ex projection, k em,filtered -k em projection, and k em -k exProjection) to determine if the power range defined at 1402 is sufficient. In one aspect, the error surface can be examined to confirm that the mapping includes a well-defined minimum. In this aspect, if the examination of the error mapping does not identify a minimum, the range of one or more power values can be modified and method 1400 can be repeated. In one aspect, an assessment of the optical properties of the patient's skin (e.g., melanin absorption, blood content, and / or scattering coefficient) can be used to classify the patient, and thus an appropriate set of coefficients can be selected for that class.

[0299] In one aspect, the exponent k can be stored ex , k em and k em,filtered in combination and used for subsequent measurements performed by system 200. Figure 18 is a graph comparing the raw fluorescence signal (blue line) with the calculated IF signal (red line) for the measurement dataset. In another aspect, global correction can be calculated by combining the measurements obtained using multiple different systems and / or sensor heads and identifying the combination of exponents corresponding to the overall minimum error value.

[0300] Linear regression method

[0301] In one aspect, the value of the power coefficient used in equation (20) above is empirically determined using a linear regression method. A flowchart showing the various steps of the linear regression method that obtains a correction in the form of a regression equation with predictor variables (DR Figure 19 is provided in. ex , DR em , DR em,filtered ).

[0302] Referring to Figure 19 , method 1900 can include logarithmically transforming Flr to log(Flr) to prepare the raw fluorescence measurement Flr for analysis at 1902. Figure 20 is a graph of log(Flr) generated at 1902. Referring again to Figure 19 , in one aspect, method 1900 can also include selecting a region of stable optical properties 2002 (see Figure 20 ), and in this aspect, the region of stable optical properties 2002 generally corresponds to a linear segment on the log(Flr) graph as shown in Figure 20 . In this aspect, method 1900 also includes obtaining a linear regression model 2004 within the region of stable optical properties 2002 at 1906. The linear regression model 2004 can be obtained using any regression method without limitation, including but not limited to multivariate linear regression modeling methods.

[0303] Referring againFigure 19 Method 1900 may also include extending, at 1908, a linear regression model 2004 obtained within a region of stable optical property 2002 to produce an extended linear regression 2008 that extends into a region of variable optical property 2010. In one aspect, the region of variable optical property 2010 is characterized by a non-linear distribution within the log(Flr) curve plot as shown Figure 20 in the log(Flr) curve plot shown.

[0304] Referring again to Figure 19 , method 1900 may also include obtaining a linear regression model 2004 that has predictor variables Flr, DR ex , DR em and DR em,filtered at 1910 and linear curve fitting 2004 as the predicted response. The extension of the linear regression 2008 generated at 1908 can be used to train the linear regression model obtained at 1910.

[0305] In one aspect, a measurement data set obtained from a single individual and / or a single system and sensor head can be used to develop a linear regression model. In another aspect, multiple measurement data sets obtained from multiple individuals and / or multiple systems and sensor heads can be used to develop a linear regression model. In some aspects, the linear regression model can be re-developed for each new measurement data set obtained for an individual. In at least some other aspects, constants and parameters characterizing the linear regression model developed as described above can be stored for subsequent re-development of the linear regression model in place of each measurement data set obtained as described above.

[0306] d) Fault detection sub-unit

[0307] Referring again to Figure 13 , the processing unit 236 of the controller 212 may also include a fault detection sub-unit 1312 that is configured to monitor the functionality of the light sources 218 / 220 and the photodetectors 222 / 224 and to notify the user via the display unit 216 of any irregularities of any detected faults within the system 200. In various aspects, the fault detection sub-unit 1312 can monitor the functionality of the light sources 218 / 220 and the photodetectors 222 / 224 by examining the signal levels received from the light sources 218 / 220 and the photodetectors 222 / 224 and additional temperature sensors 228 and additional photodetectors 226 of the associated sensor head 204 (see Figure 2) to achieve basic recognition of fault and notification status. In various aspects, the signal amplitude (see Equation (1)) and the average signal can be used to determine the peak and lowest point levels of the modulation of the LED light sources 218 / 220. The lowest point of the signal, defined herein as the average signal minus half of the peak-to-peak signal, can be used in one aspect to monitor the ambient light level. Without being limited to any particular theory, for the additional contribution to the lowest point level of the modulation signal, such as amplifier DC offset, can be ignored as being small and constant relative to the contribution of ambient light leakage. In one aspect, if the ambient light level recorded at a low detector amplifier gain exceeds approximately one-quarter of the range of the high-speed ADC 1102, an ambient light notification is issued to the user via the display unit 216.

[0308] In various other aspects, saturation of the photodetectors 222 / 224 can also be monitored by the fault detection sub-unit 1312. In these other aspects, saturation can be monitored by calculating the peak of the signal, defined herein as the average signal value plus half of the peak-to-peak signal. If the peak of the signal falls within 5% of the saturation of the ADC range, the fault detection sub-unit 1312 can issue a saturation notification to the user via the display unit 216. If the fault detection sub-unit 1312 detects a saturation event, then the ambient light level can then be checked to determine whether the saturation event is associated with ambient light saturation, defined herein as a saturation event that occurs simultaneously with the ambient light notification as described above. If an ambient light saturation event is detected, the fault detection sub-unit 1312 issues an ambient light saturation notification to the user via the display unit 216, and the data acquisition performed by the acquisition unit 234 continues in this notification state to allow the user to resolve the situation. If a saturation event is detected that is not associated with excessive ambient light, the fault detection unit can signal the photodetector 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 unit issues a notification to the user via the display unit to report an ambient light saturation event or a saturation event that is not related to excessive ambient light. In some aspects, if a saturation event is detected, but when the system 200 is configured in the engineering mode as described above, the user has disabled the automatic gain control, the user is also notified via the display unit.

[0309] e) Post-reagent application selection sub-unit

[0310] Referring again to Figure 13 , the processing unit 236 may also include a post-reagent application selection sub-unit 1308, which is configured to automatically identify the portion of the measurement data set corresponding to the post-reagent application periods 1508 / 1510 (see Figure 15A ). Referring to Figure 15A, as described above, after injecting an exogenous fluorophore such as MB-102 into the bloodstream of a patient, the exogenous fluorophore undergoes an equilibration period of diffusing from the bloodstream into the rest of the extracellular tissue of the patient. After reagent injection 1503, the temporal profile of the fluorescence signal Flr can be characterized as a biexponential signal profile described by Equation (22):

[0311]

[0312] where C0 is the baseline signal that is typically removed by baseline subtraction as described above.

[0313] Referring again to Figure 15A , once the exogenous fluorophore has diffused into the extracellular tissue of the patient to reach quasi-steady state conditions, a post-equilibration period 1510 is reached, and the fluorescence signal can be characterized as a linear decay. Without being limited to any particular theory, it is assumed that the post-equilibration region of the measurement data set is characterized as the region of the IF temporal profile that, when log-transformed, is well described by a linear equation. In one aspect, the post-equilibration region is well described by Equation (23):

[0314] IF post-equilibration = C0 + C1e -t / τ Equation (23)

[0315] In one aspect, the post-reagent administration selector unit 1308 can automatically identify the post-reagent administration period 1510 by performing a monoexponential curve fit at different portions of the IF data set and analyzing the associated curve fit error in each of the different portions. In various aspects, the post-reagent administration selector unit 1308 can select the earliest-occurring portion of the IF data set where the curve fit error associated with the monoexponential curve fit is below a threshold as the initial post-reagent administration portion of the IF data set suitable for data correction and analysis as described above. Any analysis method suitable for comparing the curve fit errors associated with monoexponential curve fits of different portions of the IF data set can be used in the post-reagent administration selector unit 1308, including but not limited to linear curve fit portions of the IF data set that fall 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 selector unit 1308 can generate at least one signal that is configured to signal to the diffuse reflectance correction unit 1306 and / or the RDTC calculation unit 1310 the time range within the IF data set corresponding to the post-reagent administration period 1508 / 1510 to enable selection of an appropriate portion of the IF data set for correction and analysis as disclosed herein.

[0316] In another aspect, a linear fit and a 2-exponential fit of the IF data can be compared. In this other aspect, once the fit errors are equivalent (corrected for the additional degrees of freedom in the 2-exponential fit), the equilibrium can be identified as complete.

[0317] f) RDTC calculation subunit

[0318] In various aspects, the system 200 is configured to convert various measurements from the photodetectors 222 / 224 and associated light sources 218 / 220 and other thermal and optical sensors in response to illumination by light at the excitation wavelength into a corrected intrinsic fluorescence (IF) signal corresponding to the detected fluorescence, which corrected intrinsic fluorescence (IF) signal can be attributed solely to the fluorescence emitted by the exogenous fluorophore at the emission wavelength. In various aspects, the exponential decay of the IF signal during the post-reagent administration portion of the IF data set can be analyzed to monitor and quantify renal function.

[0319] In one aspect, the exponential decay of the IF signal during the post-reagent administration portion of the IF data set can be converted into a glomerular filtration rate (GFR), which is configured to quantify renal function. In another aspect, the exponential decay of the IF signal during the post-equilibrium portion of the IF data set can be converted into a renal decay time constant (RDTC), which is also configured to quantify renal function. In another aspect, the exponential decay of the IF signal during the post-equilibrium portion of the IF data set can be converted into a renal decay rate, which is also configured to quantify renal function.

[0320] Referring again to Figure 13 , the processing unit 236 may also 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 mono-exponential decay described in Equation (23) above. In one aspect, after precise baseline subtraction by the baseline subtraction subunit 1304, the renal decay time constant τ can be calculated by performing a linear regression on the logarithmically transformed IF signal data (log(IF)), as described in Equation (24):

[0321]

[0322] In various aspects, the RDTC calculation subunit 1310 can generate a signal that is configured to use the display unit 216 to generate a display of the calculated RDTC. The display of the calculated RDTC can be provided to the display unit 216 in any suitable form, 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 representations configured to determine whether the calculated RDTC can be classified as normal / healthy, abnormal, high, low, and any other suitable classification. In various other aspects, when additional data is obtained and analyzed, any of the above graphical forms can be updated continuously or discontinuously. In one aspect, the RDTC calculation subunit 1310 can calculate the RDTC as described above within non-overlapping and / or overlapping windows in the IF dataset.

[0323] In another aspect, the RDTC calculation subunit 1310 can use known methods to convert the RDTC into a glomerular filtration rate (GFR). In this regard, the RDTC can be inverted and multiplied by a slope to produce a prediction of cGFR, the GFR, which can be corrected for body size (e.g., body surface area or volume of distribution).

[0324] v) Memory

[0325] Referring again to Figure 2 , the controller 212 of the system 200 can also include a memory 242 that is configured to facilitate storing data in the system 200. In some embodiments, the memory 242 includes multiple storage components, such as but not limited to hard disk drives, flash memories, random access memories, and magnetic or optical disks. Optionally or additionally, the memory 242 can include a server such as a remote storage device that communicates with the controller 212. The 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 the controller 212 described above. In one embodiment, the memory 242 can be or include a computer-readable medium, such as a floppy disk device, a hard disk device, an optical disk device, or a magnetic tape device, a flash memory, or other similar solid-state memory devices, or an array of devices, including devices in a storage area network or other configurations. A computer program product can be tangibly embodied in an information carrier. The computer program product can also contain instructions that, when executed, perform one or more functions, such as those described herein. The information carrier can be a non-transitory computer or machine-readable medium, such as the memory 242 or the memory on the processor 238.

[0326] In various aspects, system 200 can record the raw measurements and processed data into a series of files. Each file can include a header file that contains information about the operator, tool, and session. Each experimental session records a set of files into a separate folder for each sensor head used in that session. The raw data files can include in-phase, quadrature, and average measurements from the detector and monitor during the active periods of the excitation and emission wavelength LEDs, as well as the gain settings of the LEDs and detector at the time of data acquisition.

[0327] In various other aspects, the processed data files can include fluorescence and diffuse reflectance measurements after amplitude calculation and correction of the monitor readings, as well as the gain settings of the LEDs and detector. The intrinsic fluorescence data files can include intrinsic fluorescence measurements resulting from diffuse reflectance correction of the raw fluorescence signal. The GFR files can include the calculated GFR as a function of time, which is classified to indicate whether post-equilibration has occurred, as well as confidence bounds. The telemetry files can include temperature and voltage measurements. The event log files can include user and auto-generated event records.

[0328] vi) GUI unit

[0329] Referring again to Figure 2 , in various aspects, controller 212 can include a GUI unit 240, which is configured to receive a plurality of signals from other units of the system, the plurality of signals encoding various measured and transformed data. Additionally, the GUI unit can be configured to generate a signal that is configured to operate display unit 216 to display data, frames, tables, and / or any other communication of information between the user and system 200.

[0330] vii) Processor

[0331] Referring again to Figure 2 , controller 212 can also include a processor 238. Processor 238 can include any type of conventional processor, microprocessor, or processing logic that interprets and executes instructions. Processor 238 can be configured to process instructions for execution within controller 212, including instructions stored in memory 242 for displaying graphical information of the GUI on an external input / output device (such as display unit 216 coupled to the high-speed interface). In other implementations, multiple processors and / or multiple buses and multiple memories and multiple types of memories can be appropriately used. Additionally, multiple controllers 212 can be connected to each device that provides a portion of the necessary operations to enable the functionality of system 200. In some embodiments, processor 238 can include an acquisition unit 234, a photodetector control unit 232, a light source control unit 230, and / or a processing unit 236.

[0332] As used herein, a processor such as processor 238 may include any programmable system that includes a system using a microcontroller, reduced instruction set circuit (RISC), application specific integrated circuit (ASIC), logic circuitry, and any other circuitry or processor capable of performing the functions described herein. The above examples are merely examples and are not intended to limit the definition and / or meaning of the term "processor" in any way.

[0333] As described herein, a computing device and a computer system include a processor and a memory. However, any processor in the computer device referred to herein may also refer to one or more processors, where the processors may be in one computing device or in multiple computing devices operating in parallel. Additionally, any memory in the computer device referred to herein may also refer to one or more memories, where the memories may be in one computing device or in multiple computing devices operating in parallel.

[0334] C. Operating Unit

[0335] The operating unit 214 may be configured to enable a user to interface with the controller 212 (e.g., visually, audibly, by touch, button press, stylus tap, etc.) to control the operation of the system 200. In some embodiments, the operating unit 214 may further be coupled to each sensor head 204 to control the operation of each sensor head 204.

[0336] D. Display Unit

[0337] Referring again to Figure 2 , the system 200 may further include a display unit 216, which is configured to enable a user to view data and control information of the system 200. The display unit 216 may also be coupled to other components of the system 200, such as the sensor heads 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 a client application) to the user. The graphical user interface may include, for example, a display of the GFR value as described above generated by the system 200, as well as the operation data of the system 200.

[0338] Exogenous labeling

[0339] Without being limited to any particular theory, highly hydrophilic and small (creatinine, molecular weight = 113) to medium-sized (inulin, molecular weight ~5500) molecules are known to be rapidly cleared from the systemic circulation by glomerular filtration. In addition to these properties, an ideal GFR agent is neither reabsorbed nor secreted by the renal tubules, has negligible binding to plasma proteins, and has very low toxicity. To design an optical probe that meets all of these requirements, a balance has been achieved between the photophysical properties and the molecular size and hydrophilicity of the fluorophore. For example, while hydrophobic cyanine dyes and indocyanine dyes optimally absorb and emit within 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 into highly hydrophilic species required for renal clearance, but due to the limited π-systems formed by these lower molecular weight compounds, they are typically able to excite and emit single photons in the ultraviolet (UV).

[0340] To address the pharmacokinetic issues associated with enhancing photophysical properties, simple derivatives of 2,5-diaminopyrazine-3,6-dicarboxylic acid are 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 the pharmacokinetics and photophysical properties of GFR, SAR studies have been conducted using amide-linked variants of these derivatives. A variety of hydrophilic functions can be employed to achieve rapid renal clearance of this class of pyrazine fluorophores, which include carbohydrates, alcohols, amino acids, and various PEG-based related strategies. PEG surrogates can be used to increase hydrophilicity and solubility, reduce toxicity, and modulate the aggregation of the resulting pyrazine derivatives. Variations in the molecular weight and structure (and thus hydrodynamic volume) of a series of medium-sized PEG pyrazine derivatives can also be suitable for use as endogenous fluorophores.

[0341] In one aspect, the exogenous fluorophore is MB-102.

[0342] Examples

[0343] The following examples illustrate various aspects of the disclosed systems and methods.

[0344] Example 1: Perturbation analysis

[0345] To demonstrate the effectiveness of the diffuse reflectance data correction method described above herein, the following experiment was conducted.

[0346] Using the method described above herein, particularly the diffuse reflectance data correction method, a system similar to system 200 described above herein was used to monitor the fluorescence generated during the renal elimination of the exogenous fluorophore MB-102.

[0347] Figure 21Ais a graph summarizing the change in the amplitude of the raw fluorescence signal (Flr) before and approximately 6 hours after injecting the MB-102 fluorophore into a pig. During the post-equilibration portion, corresponding to the time of approximately 13:45 in Figure 21A a pig was subjected to a selected series of perturbations to alter the optical properties of the pig's skin and / or underlying tissue: administration of a blood pressure drug to induce vasodilation / vasoconstriction 2102, application of pressure to compress the tissue 2104, lateral movement of the sensor head 2106, SpO2 reduction 2108, SpO2 reduction 2110, removal / replacement of the sensor head 2112 / 2114, and skin cooling 2116.

[0348] Figure 21B is a graph summarizing the corrected intrinsic fluorescence signal (IF) as corrected above without baseline subtraction. At time points after 2 hours post-reagent injection, the time course of the IF signal is characterized by an expected signal mono-exponential decay with decay variations due to the applied perturbations.

[0349] Table 4 summarizes the specific effects of the diffuse reflectance data correction method on the Flr data associated with each individual perturbation:

[0350] Table 4. Effect of diffuse reflectance data correction

[0351]

[0352]

[0353] Figure 21C is a graph summarizing the detected diffuse reflectance signals DR em,filtered 、DR em and DR ex substituted into equation (20) to determine the diffuse reflectance correction of the raw Flr signal as described above. As Figure 21C shown, the DR em,filtered signal is the most sensitive to the various perturbations. The DR em and DR ex signals exhibit a moderate change in response to the perturbations.

[0354] The results of these experiments demonstrate that diffuse reflectance data correction is capable of correcting the raw fluorescence signal data to compensate for the effects of various perturbations that cause various changes in the optical properties of the skin.

[0355] Example 2: Sensor head with flared housing

[0356] Figure 23is a perspective view of the sensor head 204a in another aspect. In this other aspect, the sensor head 204a includes a housing 600a formed by an upper housing 602a and a flared lower housing 604a. The surface area of the lower housing 604a is enlarged to form an enlarged bottom surface 608a. The housing 600a also includes a cable opening 806a formed through the upper housing 602a.

[0357] Figure 24 is a bottom view of the sensor head 204a showing the bottom surface 608a of the housing 600a. The bottom surface 608a may include a perforated plate 702a that includes one or more holes 704a configured to transmit light between the patient's skin and a light source and a light detector contained within the housing 600. As Figure 24 shown, the holes 704a include a light transmission hole 1002a configured to transmit illumination generated by the first light source 218 / 220 and the second light source 218 / 220 to the tissue of the patient 202, and first and second detector holes 1004 / 1006 configured to receive light from the tissue of the patient 202. In one aspect, the bottom surface 608a is capable of positioning the holes 704a beneath a relatively large area that is shielded from ambient light conditions by the bottom surface 608a. This reduction in scattered ambient light entering the first detector holes 1004 / 1006 and the second detector holes 1004 / 1006 reduces the noise introduced into the light intensity measurements obtained by the first and second light detectors 222 / 224.

[0358] In various aspects, the bottom surface 608a of the housing 600a may be attached to the patient's skin using a biocompatible and transparent adhesive material 610a, which includes but is not limited to a transparent double-sided medical-grade adhesive, as Figure 24 shown. The transparent adhesive material 610a may be positioned on the bottom surface 608a such that the adhesive material 610a covers the holes 704a.

[0359] 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 is Figure 25 an exploded view of the inner housing 2502 and associated electrical components shown in Figure 25 and Figure 26, the inner housing 2502 is contained within the housing 600a and is mounted to the lower housing 608a. The 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 photodetector 222 is mounted within the first detection well 908, and a second photodetector 224 is mounted within the second detection well 910. First and second light sources 218 / 220 are mounted within the light source well 902. In one aspect, the first detection well 908, the second detection well 910, and the light source well 902 of the sensor mount 912 are optically isolated from each other to ensure that light from the light sources 218 / 220 does not reach the photodetectors 222 / 224 without being coupled through the skin of the patient 202. As detailed above, the separation between the two detection wells 908 / 910 ensures that the detected fluorescence signal from the exogenous fluorophore is distinguishable from the unfiltered excitation light.

[0360] Referring Figure 26 , the inner housing 2502 includes a first detection aperture 2602, a second detection aperture 2604, and a light source aperture 2606. The sensor mount 912 is coupled to the inner housing 2502 such that the first detection aperture 2602, the second detection aperture 2604, and the light source aperture 2606 are aligned with the first detection well 908, the second detection well 910, and the light source well 902 of the sensor mount 912, respectively.

[0361] 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. Additionally, diffusers 2616, 2618, and 2620 are coupled to the optically transparent windows 2610, 2612, and 2614, respectively. The diffusers 2616, 2618, and 2620 are arranged to spatially homogenize the light transmitted from the light sources 218 / 220 to the tissue and to spatially homogenize the light detected by the photodetectors 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.

[0362] In view of the foregoing, it can be seen that several advantages of the present disclosure are achieved and other favorable results are obtained. Since various changes may be made to the above methods and systems without departing from the scope of the present disclosure, all of the content included in the above description and shown in the accompanying drawings should be construed in an illustrative rather than a limiting sense.

[0363] When introducing an element or various versions, embodiments or aspects of the present disclosure, the articles "a", "an", "the" and "said" are intended to mean that there is one or more elements. The terms "comprising", "including" and "having" are intended to be inclusive and mean that there may be additional elements in addition to the listed elements.

Claims

1. A sensor head, comprising: a housing configured as an upper housing and a lower housing, wherein a bottom surface of the lower housing includes a contact surface configured to be attached to a patient's skin; at least two light sources configured to transmit light into a first region of the patient, the at least two light sources including a first blue LED light source and a second green LED light source, wherein the first blue LED light source transmits light in an excitation spectrum including an excitation wavelength, and the second green LED light source transmits light in an emission spectrum including an emission wavelength; and at least one light detector including a filter that overlaps with the excitation spectrum and the emission spectrum and is configured to block light at the excitation wavelength, wherein the at least one light detector is configured to detect fluorescence including the emission wavelength at a third region of the patient; and at least one unfiltered light detector configured to detect fluorescence including the excitation wavelength and the emission wavelength at a second region of the patient; wherein the housing encloses the at least two light sources, the at least one light detector, and the at least one unfiltered light detector.

2. The sensor head according to claim 1, wherein, The contact surface is configured to be attached to the patient's skin using a biocompatible adhesive material.

3. The sensor head according to claim 2, wherein, The adhesive material is configured to transmit light through the at least two light sources into the patient and from the patient to the light detector.

4. The sensor head according to claim 2, wherein, The adhesive material is an optically transparent material.

5. The sensor head according to claim 2, wherein, The adhesive material is made of a non-fluorescent material.

6. The sensor head according to claim 1, wherein, The contact surface includes a perforated plate having one or more holes configured to transmit light between the patient's skin and the at least two light sources and the at least one light detector.

7. The sensor head according to claim 1, wherein, The contact surface includes a temperature sensor opening configured to provide a thermal path from the patient's skin surface to a temperature sensor inside the housing.

8. The sensor head according to claim 6, wherein, The perforated plate includes a sensor mount configured to inhibit light leakage between the at least one light source and the at least one light detector.

9. The sensor head according to claim 8, wherein, The sensor mount is made of a conductive material.

10. The sensor head according to claim 6, wherein, The holes are positioned toward the center of the contact surface.

11. The sensor head according to claim 6, wherein, The bottom surface of the housing is configured to be flared to provide a larger surface area for contacting the patient's body surface.

12. The sensor head according to claim 11, wherein, The contact surface is configured to be attached to the patient's skin using a transparent adhesive material.

13. The sensor head according to claim 12, wherein, The transparent adhesive material is positioned on the bottom surface such that the adhesive material covers the holes.

14. A sensor head, comprising a housing, the housing being configured as an upper housing and a lower housing, wherein, The bottom surface of the lower housing includes a contact surface configured to be attached to the patient's skin, and the housing includes: (i) an inner housing having a sensor mount; (ii) at least two light sources configured to transmit light into a first region of the patient, wherein at least one light source transmits light in an excitation spectrum including an excitation wavelength, and wherein at least one light source transmits light in an emission spectrum including an emission wavelength; and (iii) at least one light detector including a filter that overlaps with the excitation spectrum and the emission spectrum and is configured to block light at the excitation wavelength, wherein the at least one light detector is configured to detect fluorescence including the emission wavelength at a third region of the patient; and at least one non-filtered light detector configured to detect fluorescence including the excitation wavelength and the emission wavelength at a second region of the patient, wherein the housing encloses the at least two light sources and the at least two light detectors, wherein the sensor mount includes: a first detection well having a first light detector mounted therein; a second detection well having a second light detector mounted therein; and a light source well having a first light source and a second light source mounted therein, and wherein the first detection well, the second detection well, and the light source well are configured to be optically isolated from each other to ensure that light from the light source does not reach the light detector without being coupled through the body surface of the patient.

15. The sensor head according to claim 14, wherein, The inner housing further includes: a first detection hole configured to be aligned with the first detection well; a second detection hole configured to be aligned with the second detection well; and a light source hole configured to be aligned with the light source well.

16. The sensor head according to claim 15, wherein, The inner housing further includes an optically transparent window configured to be coupled within the first detection hole, the second detection hole, and the light source hole.

17. The sensor head according to claim 16, wherein, The inner housing further includes a diffuser configured to be coupled to the optically transparent window.

18. The sensor head according to claim 14, wherein, The first light source is a first blue LED light source including a blue LED, and the second light source is a second green LED light source including a green LED, and wherein the first blue LED light source transmits light at the excitation wavelength, and wherein the second green LED light source transmits light at the emission wavelength.

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