Non-invasive monitoring method using diffuse reflectance corrected fluorescent tracer

Abstract

A non-invasive monitoring method of a fluorescent tracer with diffuse reflectance correction is provided, the method comprising providing a measurement data set comprising a plurality of measurement entries, the measurement entries comprising at least two measurements obtained from a patient before and after administration of a fluorescent agent. The measurements can comprise one or more of: a DR ex signal detected by an unfiltered light detector during illumination by excitation wavelength light from a first region of a proximal diffuse reflecting medium; a Flr signal detected by a filtered light detector during illumination by excitation wavelength light; and a DR em signal detected by an unfiltered light detector during illumination by emission wavelength light. The method further comprises identifying a post-agent administration portion of the measurement data set; and converting each Flr signal to an IF signal representing detected fluorescence intensity emitted only by the fluorescent agent.

Description

[0001] This application is a continuation-in-part of Patent Application No. 201880008812.7, filed July 26, 2019, which is an international application of International Application No. PCT / US2018 / 016041, filed January 30, 2018, entitled "Non-invasive Monitoring Method of Fluorescent Tracers with Diffuse Reflectance Correction," the entire contents of which are incorporated herein by reference.

[0002] Related Applications

[0003] This application claims the benefit of U.S. Provisional Application No. 62 / 452,025, filed January 30, 2017, the entire contents of which are incorporated herein. TECHNICAL FIELD

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

[0005] To minimize the risk of acute renal failure due to various clinical, physiological, and pathological conditions, real-time dynamic monitoring of a patient's renal function at the bedside is highly desirable. This is particularly important in the case of critically ill or injured patients, as a significant proportion of these patients are at risk of multiple organ failure (MOF) due to one or more severe functional impairments, such as: acute lung injury (ALI), adult respiratory distress syndrome (ARDS), hypermetabolism, hypotension, persistent inflammation, and / or sepsis. Renal function can also be compromised due to renal impairment associated with the administration of nephrotoxic drugs as part of a procedure, such as angiography, diabetes, autoimmune diseases, and other functional impairments and / or injuries associated with renal damage. To assess the patient's status and monitor the severity and / or progression of renal function over time, there is considerable interest in developing a simple, accurate, and continuous method to determine 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 conjunction with patient demographic factors such as body weight, age, and / or race to estimate glomerular filtration rate (GFR), a measure of kidney function. However, creatinine-based assessments of renal function can be prone to inaccuracies due to a number of potential factors, including: age, hydration status, renal perfusion, muscle mass, dietary intake, and a number of other anthropometric and clinical variables. To compensate for these differences, a series of creatinine-based equations (recently expanded to include cystatin C) have been established that incorporate factors such as gender, race, and other relevant factors for estimating glomerular filtration rate (eGFR) based on serum creatinine measurements. However, these eGFR equations do not provide any way to compensate for the majority of the aforementioned sources of variability, and thus have relatively poor precision. Furthermore, the eGFR methods generally produce results that are 72 hours behind the true GFR.

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

[0008] The availability of real-time, accurate, and reproducible measurement of renal excretion rates using exogenous markers in a patient-specific but environment-agnostic context would represent a significant advance over any currently practiced method. Furthermore, a method that relies solely on the renal clearance of an exogenous chemical entity would provide a direct and continuous pharmacokinetic measurement without the need for subjective interpretation based on age, muscle mass, blood pressure, etc. BRIEF DESCRIPTION OF DRAWINGS

[0009] The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.

[0010] The present disclosure will be better understood with reference to the following detailed description when considered in conjunction with the following drawings, in which: i. FIG. 1 shows a schematic of a method for measuring renal function in accordance with the present disclosure. ii. FIG. 2 shows a schematic of a method for measuring renal function in accordance with the present disclosure. iii. FIG. 3 shows a schematic of a method for measuring renal function in accordance with the present disclosure. iv. FIG. 4 shows a schematic of a method for measuring renal function in accordance with the present disclosure. v. FIG. 5 shows a schematic of a method for measuring renal function in accordance with the present disclosure. vi. FIG. 6 shows a schematic of a method for measuring renal function in accordance with the present disclosure. vii. FIG. 7 shows a schematic of a method for measuring renal function in accordance with the present disclosure. viii. FIG. 8 shows a schematic of a method for measuring renal function in accordance with the present disclosure. ix. FIG. 9 shows a schematic of a method for measuring renal function in accordance with the present disclosure. x. FIG. 10 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xi. FIG. 11 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xii. FIG. 12 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xiii. FIG. 13 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xiv. FIG. 14 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xv. FIG. 15 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xvi. FIG. 16 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xvii. FIG. 17 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xviii. FIG. 18 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xix. FIG. 19 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xx. FIG. 20 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxi. FIG. 21 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxii. FIG. 22 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxiii. FIG. 23 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxiv. FIG. 24 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxv. FIG. 25 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxvi. FIG. 26 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxvii. FIG. 27 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxviii. FIG. 28 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxix. FIG. 29 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxx. FIG. 30 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxi. FIG. 31 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxii. FIG. 32 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxiii. FIG. 33 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxiv. FIG. 34 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxv. FIG. 35 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxvi. FIG. 36 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxvii. FIG. 37 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxviii. FIG. 38 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xxxix. FIG. 39 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xl. FIG. 40 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xli. FIG. 41 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlii. FIG. 42 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xliii. FIG. 43 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xliv. FIG. 44 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlv. FIG. 45 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlvi. FIG. 46 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlvii. FIG. 47 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlviii. FIG. 48 shows a schematic of a method for measuring renal function in accordance with the present disclosure. xlix. FIG. 49 shows a schematic of a method for measuring renal function in accordance with the present disclosure. l. FIG. 50 shows a schematic of a method for measuring renal function in accordance with the present disclosure. li. FIG. 51 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lii. FIG. 52 shows a schematic of a method for measuring renal function in accordance with the present disclosure. liii. FIG. 53 shows a schematic of a method for measuring renal function in accordance with the present disclosure. liv. FIG. 54 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lv. FIG. 55 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lvi. FIG. 56 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lvii. FIG. 57 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lviii. FIG. 58 shows a schematic of a method for measuring renal function in accordance with the present disclosure. lix.

[0011] Figure 1 is a schematic diagram of a single-wavelength renal monitoring device in an aspect;

[0012] Figure 2 is a schematic diagram of a dual-wavelength renal monitoring system in an aspect;

[0013] Figure 3 is 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 over a range of optical wavelengths from about 430 nm to about 650 nm in vivo;

[0014] Figure 4 is a graph summarizing the absorption spectra of oxygenated hemoglobin (HbO2) and deoxygenated hemoglobin (Hb) defined over a range of optical wavelengths from about 200 nm to about 650 nm;

[0015] Figure 5 is a schematic diagram of the timing of light pulse cycles associated with data acquisition of a dual-wavelength renal monitoring system in an aspect, wherein each light pulse cycle includes light pulses generated at an excitation wavelength and an emission wavelength in sequence;

[0016] Figure 6 is a side view of a sensor head of a renal function monitoring system in an aspect;

[0017] Figure 7 is a bottom view of the sensor head of Figure 6 .

[0018] Figure 8 is an internal top view of the sensor head of Figure 6 , Figure 8 illustrates the arrangement of various electrical components within a housing of a sensor head of a renal function monitoring system in an aspect;

[0019] Figure 9 is an enlarged view of the internal view of Figure 8 .

[0020] Figure 10 is a schematic diagram of a hole formed within a contact face of a sensor head of a renal function monitoring system in an aspect;

[0021] Figure 11 is a schematic diagram of the simultaneous detection of light by a light detector of a sensor head in an aspect;

[0022] Figure 12 is a schematic diagram of the modulation and demodulation of light signals by a sensor head in an aspect;

[0023] Figure 13 is a block diagram illustrating sub-units of a processing unit;

[0024] Figure 14A is a flow chart illustrating the steps of a global error mapping method for determining parameters of a diffuse reflection correction equation in one aspect;

[0025] Figure 14B is a flow chart illustrating the steps of a global error mapping method for determining parameters of a diffuse reflection correction equation in a second aspect;

[0026] Figure 15A The fluorescence detected by the renal monitoring device before and after the injection of exogenous fluorescent agent (IF agent ). The data subset selected for analysis to determine the correction factor is highlighted in orange.

[0027] Figure 15B The fluorescence detected by the renal monitoring device before and after the injection of exogenous fluorescent agent (IF agent ) is a graph showing a representative intrinsic fluorescence measurement of agent Fitting to plasma derivative IF agent The data subset selected for analysis to determine the correction factors is highlighted in orange.

[0028] Figure 16 is a comparison of the logarithmic transformation of the corrected fluorescence signal measurements (log[Fit], black dashed line) and the Figure 15A and Figure 15B The selected analysis area is based on Figure 15A and Figure 15B Corrected fluorescence signal measurement (IF agent , red line).

[0029] Figure 17 is a map of representative error surfaces summarizing the normalized root mean square error (RMSE, color scale), which is the error surface for the correction factor K. ex and K em,filtered The RSME region is calculated as the difference between the linear fit of the logarithm of fluorescence and the corrected fluorescence signal measurement over a range of , where the minimum RSME region is identified by the white arrow overlaid on the map;

[0030] Figure 18 is a graph comparing raw (F, blue line) and corrected (IF, red line) fluorescence signal measurements obtained before and after injection of an exogenous fluorescent agent.

[0031] Figure 19 is a flow chart outlining the steps of a linear regression model method for determining parameters of a diffuse reflectance correction equation in one aspect;

[0032] Figure 20 is a plot of the raw fluorescence signal (Log(Flr)) that is logarithmically transformed, showing the regions of data that were used as projections suitable for: a linear regression model used to develop a data correction algorithm (orange line), a response variable for the linear regression model (black dashed line), and regions of data that were significantly varying for training the linear regression model (blue line);

[0033] Figure 21A is a plot of the raw fluorescence signal measurements taken before and after the injection of exogenous fluorescent agent. Measurements of the raw fluorescence signal were taken during exposure to various perturbations represented as colored regions beginning at approximately 13:50 hours. The various perturbations included changes in blood oxygen in the test subject, application and removal of pressure to the measured region, administration of blood pressure medication to the test subject, cooling of the measured region, and removal / replacement of the sensor head of the device;

[0034] Figure 21B is a plot of the corrected fluorescence signal measurements of Figure 21A ;

[0035] Figure 21C is a plot of the diffuse reflectance signal measurements taken simultaneously with the raw fluorescence signal measurements of Figure 21A . These signals were used to correct the raw fluorescence signal measurements of Figure 21A to produce the corrected signal shown in Figure 21B ;

[0036] Figure 22A is a block diagram showing multiple modules of a pre-processing subunit of an aspect;

[0037] Figure 22B is a block diagram showing multiple modules of a pre-processing subunit of a second aspect;

[0038] Figure 23 is an isometric view of a sensor head of a renal function monitoring system of a second aspect;

[0039] Figure 24 is a bottom view of the sensor head of the renal function monitoring system shown in Figure 23 ;

[0040] Figure 25 is an isometric view of the sensor head of the renal function monitoring system shown in Figure 23 , with the upper housing and various electrical components removed to expose the inner housing; and

[0041] Figure 26 is an exploded view of the inner housing of the sensor head shown in Figure 25 .

[0042] This written description uses examples to disclose the invention, including the best mode, and also to enable any person skilled in the art to practice the invention, including making and using any devices or systems and performing any incorporated methods. The patentable scope of the present invention is defined by the claims, and can include other examples that occur to those skilled in the art. Such other examples are intended to fall within the scope of the claims if they have structural elements in common with the ones specified in the claims, or if they do not differ from the ones specified in the claims materially. DETAILED DESCRIPTION

[0043] Unless defined otherwise, 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.

[0044] As used herein, a sample refers to a single discrete data value acquired from a signal and / or telemetry analog-to-digital converter (ADC) for a single acquisition / telemetry channel.

[0045] As used herein, a measurement refers to a single discrete data value created by demodulating or accumulating a sequence of samples from one acquisition channel.

[0046] As used herein, a measurement refers to a collection including demodulated in-phase, demodulated quadrature, and average measurements from one acquisition channel.

[0047] As used herein, a measurement subset refers to a collection including all measurements for all acquisition channels during illumination of a single source LED. For example, all measurements for an acquisition channel can include demodulated in-phase, demodulated quadrature, and average measurements.

[0048] As used herein, a measurement set refers to a collection including one measurement subset for each source LED.

[0049] As used herein, an acquisition refers to the entire process of acquiring a measurement set.

[0050] As used herein, a measurement sequence refers to a sequence of one or more measurement sets.

[0051] As used herein, a telemetry value refers to a single discrete data value acquired from a single channel of a telemetry ADC.

[0052] As used herein, a telemetry set refers to a collection including one telemetry value from each telemetry channel.

[0053] Figure 1is a schematic illustration of the system 100 provided as a non-limiting example in which a light detector 110 configured to detect only those photons having an emission wavelength (λ em ) is used to detect fluorescent light 102 having an emission wavelength (λ em ) from a region of interest of a patient 104. Typically, the exogenous fluorescent agent 112 generates the fluorescent light 102 in response to an excitation event including, but not limited to, illumination by light 106 at an excitation wavelength (λ ex ), occurrence of an enzymatic reaction, change in local electric potential, and any other known excitation event associated with the exogenous fluorescent agent. In one aspect, the system 100 can include a light source 108 configured to deliver light 106 at the excitation wavelength (λ ex ) to the patient 104. In this aspect, the fluorescent light 102 is generated in response to illumination by the light 106. Moreover, the excitation wavelength (λ ex ) of the light 106 and the emission wavelength (λ em ) of the fluorescent light 102 are spectrally distinct (i.e., λ ex is not the same as λ em ) such that the light detector 110 can be configured to selectively detect only the fluorescent light 102 by including any known light wavelength separation device including, but not limited to, an optical filter.

[0054] In some aspects, changes in the fluorescent light 102 can be monitored to obtain information about a physiological function or state of the patient. As a non-limiting example, a time-dependent decrease in the fluorescent light 102 measured after introduction of the exogenous fluorescent agent 112 into the patient's 104 circulating vasculature can be analyzed to obtain information about the renal function of the patient 104. In this non-limiting example, it can be assumed that the rate of decrease of the fluorescent light 102 is proportional to the rate at which the patient's 104 kidneys remove the exogenous fluorescent agent 112, thereby providing a measure of renal function including, but not limited to, a renal decay time constant (RDTC) and a glomerular filtration rate (GFR).

[0055] Without being bound to any particular theory, the intensity of the fluorescent light 102 detected by the light detector 110 can be affected by any one or more of a variety of factors including, but not limited to, the intensity or power of the light 106 delivered to the patient 104 at λ ex , scattering and absorption of the light 106 by intervening tissue 114 of the patient 104 between the light source 108 and the exogenous fluorescent agent 112, concentration of the exogenous fluorescent agent 112 illuminated by the light 106, and scattering and absorption of the fluorescent light 102 by intervening tissue 114 of the patient 104 between the exogenous fluorescent agent 112 and the light detector 110 at λ em .

[0056] Existing methods generally assume that the optical properties within the intervening tissue 114 remain substantially constant throughout the period over which the system 100 obtains measurements. As a result, existing methods generally obtain initial measurements through the intervening tissue 114 of the patient 104 prior to the introduction of the exogenous fluorescent agent 112, and subtract these initial measurements to correct all subsequent data obtained after the introduction of the exogenous fluorescent agent 112. However, during long-term monitoring of the patient 104, changes in the optical properties of the intervening tissue 114 may occur due to changes in at least one characteristic, including but not limited to: the efficiency of light coupling from the photodetector 110 to the patient 104; the concentration of chromophores, such as hemoglobin, due to changes in blood volume caused by vascular dilation, constriction, or compression; changes in the optical properties of chromophores, such as hemoglobin, due to changes in oxygenation state; and changes in tissue structure, such as changes associated with edema.

[0057] These dynamic changes in the optical properties of intervening tissue 114 can introduce uncertainty into long-term measurements of fluorescence 102. As a non-limiting example, changes in the optical properties of intervening tissue 114 can modulate the intensity or power of light 106 illuminating exogenous fluorescing agent 112, resulting in a modulation of fluorescence 102 generated by exogenous fluorescing agent 112, which can be erroneously interpreted as a modulation of the concentration of exogenous fluorescing agent 112. As another non-limiting example, changes in the optical properties of intervening tissue 114 can modulate the intensity or power of fluorescence 102 reaching light detector 110, which can also be erroneously interpreted as a modulation of the concentration of exogenous fluorescing agent 112. The potential modulation of changes in the optical properties of intervening tissue 114 can introduce uncertainty into measurements of fluorescence 102, particularly those associated with long-term monitoring of fluorescence 102, as described above.

[0058] In various aspects, a method is provided for correcting in vivo, real-time measurements of fluorescence from an exogenous fluorescent agent to eliminate the effects of variations in optical properties within a patient's tissue. Including an additional measurement of light passing through the patient's tissue via an optical path separate from that of the fluorescence measurement (i.e., diffuse reflection) enhances the quantification of changes in the optical properties of the tissue during prolonged monitoring of fluorescence from the exogenous fluorescent agent within the patient. Including this additional measurement in the correction method, in various aspects, has been found to significantly enhance the fidelity of the fluorescence measurement, even in the presence of significant perturbations, as described below.

[0059] A detailed description of an apparatus for monitoring the fluorescence of an exogenous fluorescent agent in vivo and a method for correcting fluorescence measurements to eliminate the effects of diffuse reflection of light within patient tissue is provided below.

[0060] While the devices and methods are described below in the context of a non-invasive optical renal function monitor, it should be understood that the correction methods described herein can be applied to any compatible device configured to perform measurements by transmitting EM radiation from an external source through any scattering medium and / or receiving EM radiation propagating through any scattering medium to an external detector, with appropriate modifications. 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 of at least one EM frequency without limitation. At least a portion of the scattering medium can also include one or more sub-structures or compounds capable of reflecting and / or absorbing EM radiation. Non-limiting examples of scattering media include tissue of a living or dead organism, such as the skin of a mammal; a gas, such as air with or without additional particulates, such as dust, liquid droplets, or solid particulate material; a fluid, such as water with or without additional particulates, such as air bubbles or solid particulate material. Furthermore, the devices and methods described below are not limited to the detection of renal function, but can be modified for use in detecting the function of other physiological systems, including but not limited to the liver system or the gastrointestinal system.

[0061] System description

[0062] In various aspects, the methods of correcting fluorescence measurements to eliminate the effects of local skin property variations can be incorporated into any fluorescence monitoring system, including but not limited to systems for monitoring renal function in vivo in real time and optically by measuring changes in fluorescence of an exogenous fluorescent agent injected into a patient as the agent is re-eliminated from the patient's body. Figure 2 is a block diagram of a system 200 for optically monitoring renal function of a patient 202 in one aspect through fluorescence measurements of an exogenous fluorescent agent injected in 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.

[0063] The system 200 can also include a controller 212, an operating unit 214, and a display unit 216 operably coupled to the at least one sensor head 204. In various aspects, the controller 212 is configured to control the operation of the at least one sensor head 204, as described in further detail below. The controller 212 is also configured to receive measurements of light from the at least one sensor head 204. The controller 212 is further configured to correct the measurements of light corresponding to fluorescence from an exogenous fluorescent agent according to at least one method, including but not limited to the disclosed methods of using diffuse reflection of light to correct fluorescence measurements. The controller 212 is also configured to convert the fluorescence measurements received from the at least one sensor head 204 into summary parameters representative of renal function of the patient 202. In addition, the controller 212 is configured to receive at least one signal representative of user input from the operating 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).

[0064] Detailed descriptions of the sensor head 204 and the controller 212 are provided below.

[0065] A. Sensor Head

[0066] 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 includes an upper housing 602 and a lower housing 604 connected together to enclose two light sources and two light detectors. The bottom surface 608 of the lower housing 604 also includes a contact surface 606 configured to be attached to the skin of the patient 202 using a biocompatible adhesive material, including but not limited to a surgical adhesive. In use, the surface of the adhesive material opposite the contact surface 606 can be affixed to the skin of the patient 202. In various aspects, the adhesive material can be configured to transmit light through the light sources into the patient and also to transmit fluorescence from the patient to the light detectors. In one aspect, the adhesive material can be an optically transparent material. In another aspect, the adhesive material can be made of a non-fluorescent material to prevent the adhesive material from producing spurious fluorescence.

[0067] In various other aspects, the upper housing 602 can also include one or more openings 806 configured to provide access to the interior of a cable, including but not limited to a USB cable, and / or to provide a window for displays, such as indicator LEDs, generated by circuitry contained within the housing 600.

[0068] Figure 7 is Figure 8A bottom view of the housing 600 is shown. The contact face 606 can include an aperture plate 702 that includes one or more apertures 704 configured to transmit light between the patient's skin and the light source and light detector housed within the housing 600. In one aspect, the aperture plate 702 can be epoxied into the lower housing 604 to prevent liquid ingress into the interior of the housing 600. In various aspects, the size, arrangement, and / or spacing of the one or more apertures 704 can be selected to enhance various aspects of the operation of the system 200, as described in further detail below. In another aspect, the contact face 606 can also include a temperature sensor opening 706 configured to provide a thermal path from the patient's skin surface to an additional temperature sensor 228 configured to monitor the temperature at the patient's skin surface.

[0069] Figure 8 is a schematic diagram showing the arrangement of electrical components within the housing 600. With reference to Figure 8 , the upper housing 602 and the lower housing 604 can be secured together with screws 802, and the interface between the screw holes and the two housing pieces can be filled with a waterproofing filler material 804 including, but not limited to, a silicone material such as room temperature vulcanizing silicone (RTV) to prevent liquid ingress into the interior of the housing 600.

[0070] In one aspect, the housing 600 can also include a cable opening 806 formed through the upper housing 602. The cable opening 806 can be configured to provide access to the interior of an electrical cable including, but not limited to, a USB cable. In one aspect, the cable can be capable of powering the light source, the light detector, the indicator lights, and the associated electrical devices and circuitry, as described below. In another aspect, the cable can also enable control signals to be transmitted into the housing to enable the electrical components within the housing 600 to operate, and the cable can also enable data signals to be transmitted that encode measurements obtained by one or more of the sensor devices contained within the housing 600 including, but not limited to: the first light detector 222, the second light detector 224, any additional light detectors such as the first monitor photodiode 904 and the second monitor diode 906, and any additional temperature sensors 228 (see Figure 9 ). In one aspect, the cable can be attached to the cable opening 806 and the adjacent upper housing 602 with a light absorbing adhesive including, but not limited to, black epoxy, and can also be sealed with a waterproofing filler material including, but not limited to, RTV to prevent water intrusion.

[0071] In another aspect, the housing 600 can also include at least one display opening 808 formed through the upper housing 602. In one aspect, each display opening 808 can be configured to provide a window for a display generated by the circuitry contained within the housing 600, such as an indicator LED 810. In one aspect, each indicator LED 810 can be positioned on a circuit board 812. In one aspect, a light pipe 814 can be epoxied into the display opening 808 within the upper housing 602 above each indicator LED 810. Each light pipe 814 can be filled with a waterproof fill material, such as RTV, for liquid ingress protection. In various aspects, the at least one indicator LED 810 can illuminate in a predetermined pattern to enable a user of the system 200 to monitor the operational status of the sensor head 204.

[0072] Figure 9 is a close-up view of the interior optical region of the sensor head 204, illustrating the arrangement of the light sources 218 / 220 and the light detectors 222 / 224 within the housing 600 in one aspect. 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 that is secured to the aperture plate 702. In one aspect, the sensor mount 912 ensures that light from the light sources 218 / 220 does not reach the light detectors 222 / 224 without being coupled through the skin of the patient 202. The separation between the first light detector 222 within the first detection well 908 and the second light detector 224 within the second detection well 910 ensures that the fluorescent signal generated by the exogenous fluorescent agent within the tissue of the patient 202 is distinguishable from the unfiltered excitation light introduced by the first light source 218.

[0073] Referring again to Figure 9 The sensor mount 912 can be aligned to a circuit board (not shown) containing the light sources 218 / 220 and the light detectors 222 / 224 using an alignment pin 914, and held in place using a screw 916. In one aspect, the sensor mount 912 can be affixed 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 aspect, this light resistant joint between the circuit board and the sensor mount 912 inhibits light leakage between the light sources 218 / 220 and the light detectors 222 / 224, and also inhibits light leakage between the first light detector 222 and the second light detector 224. The aperture 704, which is configured to transmit light to and from the skin beneath the contact face 606 of the sensor head 204, is through a structurally separate aperture plate 702 (see Figure 7) formed to provide precise alignment of the aperture 704 with the corresponding light source 218 / 220 and light detector 222 / 224, described in greater detail below.

[0074] 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 an electrically conductive material, including but not limited to aluminum and aluminum alloys. In this aspect, the sensor mount 912 can be electrically coupled to the ground of the circuit board using electrically conductive screws 916. In addition, any glass windows located within the source well 902 and / or detector wells 908 / 910 adjacent to the aperture plate 702, including but not limited to the optical filter 244 and the clear glass 246 (see Figure 2 ), described below, can further include an electrically conductive coating. Non-limiting examples of suitable electrically conductive coatings for the glass windows of the sensor mount include electrically conductive indium tin oxide (ITO) coatings and any other suitable transparent and conductive coatings.

[0075] Without being limited to any particular theory, the electrically 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 be completed on the glass windows within the source well 902 and / or detector wells 908 / 910 with an electrically conductive ITO coating. In one aspect, the electrically conductive coating on the glass windows, such as an 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 electrically conductive sensor mount 912 using any known method of electrical grounding, including but not limited to connecting the glass coating to a wire of the sensor mount 912 that is attached at both ends of the wire with electrically 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 electrically conductive epoxy.

[0076] In various aspects, the contact surface 606 of the housing 600 can be attached to the skin of the patient using a biocompatible and adhesive material 610, including but not limited to a clear, double-sided medical grade adhesive, as Figure 6 and Figure 7The adhesive material 610 can be positioned on the contact face 606 such that the adhesive material covers the aperture 704, but exposes the temperature sensor opening 706 to ensure full thermal contact with the skin of the patient 202. In one aspect, the sensor head 204 can be further secured to the patient 202 using one or more additional biocompatible medical fastener devices including, but not limited to, Tegaderm bandages, medical tape, or any other suitable biocompatible medical fastener device as desired.

[0077] In one aspect, the contact face 606 can be located near the leading edge of the sensor head 204 to provide accurate positioning of the contact face 606 on a selected area of the patient's skin. In another aspect, the aperture 704 can be positioned toward the center of the contact face 606 to reduce ambient light entry. Without being limited to any particular theory, ambient light can enter one or more of the apertures 704 due to incomplete adhesion of the contact face 606 to the patient's skin and / or due to ambient light propagating into the aperture 704 through the patient's exposed skin just outside the footprint of the contact face 606.

[0078] Referring again to Figure 6 , the bottom surface 608 of the sensor head 204 is curved away from the plane of the contact face 606 to enable the sensor head 204 to be attached to different body types and locations. To attach the sensor head 204 to a relatively flat or concave surface, any gap 612 between the bottom surface 608 and the skin surface of the patient 202 can be filled with a biocompatible foam to ensure consistent contact with the patient 202.

[0079] i) light source

[0080] In various aspects, each sensor head 204 includes a first light source 218 and a second light source 220 configured to deliver light to the first region 206 of the patient 202. The first light source 218 is configured to deliver light at an excitation wavelength, and the second light source 220 is configured to deliver light at an emission wavelength. In one aspect, the excitation wavelength can be selected to fall within a spectral range in which the exogenous fluorescent agent exhibits relatively high absorption. In another aspect, the emission wavelength can be selected to fall within a spectral range in which the exogenous fluorescent agent exhibits relatively high emission. The exogenous fluorescent agent can be selected to enhance contrast relative to other chromophores within the tissue of the patient 202, including but not limited to hemoglobin within red blood cells and / or melanin within melanocytes. In various aspects, the exogenous fluorescent agent can be selected to measure within a spectral range in which other chromophores within the tissue of the patient 202, such as hemoglobin, exhibit lower changes in absorption during use.

[0081] Without being limited to any particular theory, hemoglobin (Hb) is an absorber of visible light in the tissues of the patient 202, and if the Hb absorber changes during the measurement period of the system 200, it is possible to interfere with the fluorescence measurement of the exogenous fluorescent agent. Because hemoglobin (Hb) is capable of gas exchange in nearly all tissues containing circulating blood vessels, nearly all tissues are susceptible to interference with the fluorescence measurement of the system 200 due to fluctuations in hemoglobin concentration. In most tissues, externally applied pressure can cause blood pooling, which can manifest as a significant attenuation of the fluorescence measured at the skin surface. Periodic opening and closing of blood vessels near the skin surface (“vascular motion”) can also cause fluctuations in hemoglobin concentration, which can introduce additional noise to the fluorescence measurement of the system 200 of the exogenous fluorescent agent. In addition, in some patients 202, such as those with pulmonary disease, 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 shown in Figure 3

[0082] In one aspect, the excitation wavelength and the emission wavelength for the exogenous fluorescent agent can be selected to coincide with a pair of HbO2 / Hb isosbestic points, each of which is defined herein as a wavelength at which the optical absorbance of HbO2 and Hb are approximately equal. Without being limited to any particular theory, as long as the combined concentration of HbO2 and Hb remains relatively stable during the fluorescence measurement of the system 200, the fluorescence measurement made at each isosbestic wavelength is less sensitive to changes due to changes in the oxygenation of hemoglobin. Non-limiting examples of Hb / HbO2 isosbestic wavelengths include: about 390 nm, about 422 nm, about 452 nm, about 500 nm, about 530 nm, about 538 nm, about 545 nm, about 570 nm, about 584 nm, about 617 nm, about 621 nm, about 653 nm, and about 805 nm.

[0083] In various aspects, the excitation wavelength and the emission wavelength can be selected based on the absorption wavelength and the emission wavelength of the selected exogenous fluorescent agent of the system 200. In one aspect, the excitation wavelength can be an HbO2 / Hb isosbestic wavelength and, at the same time, a wavelength within the high absorbance spectral range of the exogenous fluorescent agent. In another aspect, the emission wavelength can be an HbO2 / Hb isosbestic wavelength and, at the same time, a wavelength within the spectral range emitted by the exogenous fluorescent agent. Table 3 provides a summary of HbO2 / Hb isosbestic wavelengths within the 200 nm to about 1000 nm spectral range. Figure 4 is a plot of the absorption spectra used to identify the HbO2 / Hb isosbestic wavelengths of Table 1.

[0084] ​Table 1. HbO2 / Hb isosbestic wavelengths λ = 200 to 1000 nm

[0085]

[0086] As an illustrative example, Figure 3 is a plot summarizing the absorption spectra of HbO2and Hb, as well as the absorption and emission spectra of the exogenous fluorescent agent MB-102. The emission spectra of a blue LED light source and a green LED light source are also shown superimposed on the other spectra of Figure 3 In this aspect, the system 200 can include a blue LED as the first light source 218, and the excitation wavelength of the system 200 can be an isosbestic wavelength of about 450 nm. As listed in Table 1 and shown in Figure 3 the Hb absorbance spectrum is significantly tilted at isosbestic wavelengths of about 420 nm to about 450 nm (see columns 3 and 4 of Table 1), indicating that the relative absorbance of HbO2and Hb at an isosbestic wavelength of about 450 nm is sensitive to small changes in the excitation wavelength. However, at wavelengths higher than about 500 nm, the HbO2 / Hb spectrum is less steeply tilted, and a broader bandwidth light source (including but not limited to an LED with a bandpass filter) can be sufficient to use as the first light source 218.

[0087] In another aspect, the excitation wavelength can be selected to enhance the light absorption contrast between the exogenous fluorescent agent and the chromophore within the tissue of the patient 202. As a non-limiting example, as shown in Figure 3 at an isosbestic wavelength of 452 nm, the light absorption of MB-102 is more than three times higher than the light absorption of HbO2and Hb. Without being bound to any particular theory, MB-102 will absorb a higher proportion of the light illuminating the tissue of the patient 202 at a wavelength of about 450 nm relative to HbO2and Hb, thereby enhancing the absorption efficiency of MB-102 and reducing the intensity of the light at the excitation wavelength, which is required to elicit a detectable fluorescent signal.

[0088] In various aspects, a second isosbestic wavelength can also be selected as the emission wavelength of the system 200. As a non-limiting example, Figure 3An emission spectrum of the MB-102 exogenous contrast agent is shown, which is characterized by an emission peak at a wavelength of about 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 and second detectors 222 / 224 and 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 the chromophore within the tissue of the patient 202. Without being limited to any particular theory, the low absorbance of the chromophore at the selected emission wavelength can reduce the loss of light emitted by the exogenous fluorescent agent and increase the efficiency of the fluorescence detection.

[0089] In various aspects, the first and second light sources 218 and 220 can be any light source configured to deliver light at an excitation wavelength and an emission wavelength. Typically, the first light source 218 delivers light to the exogenous fluorescent agent at an intensity sufficient to penetrate the tissue of the patient 202, with sufficient intensity remaining to induce the exogenous fluorescent agent to emit light at the emission wavelength. Typically, the first light source 218 delivers light to the exogenous fluorescent agent at an intensity sufficient to penetrate the tissue of the patient 202, with sufficient intensity remaining after scattering and / or absorption to induce fluorescence at the emission wavelength by the exogenous fluorescent agent. However, the intensity of the light delivered by the first light source 218 is limited to an upper value to prevent adverse effects of the exogenous fluorescent agent and / or endogenous chromophores (“auto-fluorescence”) in the skin, such as tissue burning, cell damage, and / or photo-bleaching.

[0090] Similarly, the second light source 220 delivers light at the emission wavelength of the exogenous fluorescent agent at an intensity configured to provide sufficient energy to propagate through the first region 206 of the patient in a scattered and absorbed manner and with sufficient remaining intensity to propagate out of the second and third regions 208 and 210 for detection by the first and second light detectors 222 and 224, respectively. As with the first light source 218, the intensity of the light produced by the second light source 220 is limited to an upper value to prevent adverse effects such as tissue damage or photo-bleaching as previously described.

[0091] In various aspects, the first and second light sources 218 and 220 can be any light source suitable for use with a fluorescence medical imaging system and device. Non-limiting examples of suitable light sources include LEDs, diode lasers, pulsed lasers, continuously oscillating lasers, xenon arc lamps or mercury vapor lamps with excitation filters, lasers, and super-continuum light sources. In one aspect, the first and / or second light sources 218 and 220 can produce light of a narrow spectral bandwidth suitable for monitoring the concentration of an exogenous fluorescent agent using the methods described herein. In another aspect, the first and second light sources 218 and 220 can produce light of a relatively wide spectral bandwidth.

[0092] In one aspect, the selection of the intensity of the light produced by the first light source 218 and the second light source 220 of the system 200 can be influenced by any one or more of at least several factors, including but not limited to the maximum permissible exposure (MPE) for skin exposure to the laser beam according to applicable regulatory standards such as ANSI standard Z136.1. In another aspect, the light intensity for the system 200 can be selected to reduce the likelihood of photobleaching of exogenous fluorescent sources and / or other chromophores within the tissue of the patient 202, including but not limited to: collagen, keratin, elastin, hemoglobin within red blood cells, and / or melanin within melanocytes. In yet another aspect, the light intensity for the system 200 can be selected so as to elicit a detectable fluorescent signal from exogenous fluorescent sources within the tissue of the patient 202 and within the first light detector 222 and / or the second light detector. In yet another aspect, the light intensity for the system 200 can be selected to provide a suitably high light energy while reducing power consumption, inhibiting heating / overheating of the first light source 218 and the second light source 220, and / or reducing the time of patient skin exposure to light from the first light detector 222 and / or the second light detector.

[0093] 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, including but 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 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, including but not limited to: variations in the separation distance between the light sources 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 sources due to their exposure to environmental factors such as heat and humidity; variations in the sensitivity of the light detector due to its exposure to environmental factors such as heat and humidity; modulation of the duration of the illumination by the light sources, and any other relevant operable parameters.

[0094] In various aspects, the first light source 218 and the second light source 220 can be configured to modulate the intensity of the light produced as needed 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, then the intensity of the light can be electronically modulated using methods including, but not limited to, modulating the potential, current, and / or power supplied to the first light source 218 and / or the second light source 220. In another aspect, the intensity of the light can be modulated using optical methods including, but not limited to, partially or completely occluding the light exiting the first light source 218 and the second light source 220 using optical devices including, but not limited to, an iris, a shutter, and / or one or more filters; and diverting the path of the light exiting the first light source 218 and the second light source 220 using optical devices including, but not limited to, a lens, a mirror, and / or a prism from the first region 206 of the patient.

[0095] In various aspects, the intensity of the light produced 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 produced light beam. In one aspect, the laser flux can be limited to a range defined by safety standards including, but not limited to, ANSI standards for exposure to laser energy such as ANSI Z136.1. Without being limited to any particular theory, the maximum light flux delivered to the patient 202 can be influenced by a variety of factors including, but not limited to, the wavelength of the light delivered and the duration of exposure to the light. In various aspects, the maximum light flux is in the range of about 0.003 J / cm2for light delivered at wavelengths less than about 302 nm to about 1 J / cm2for light delivered in the wavelength range of about 1500 nm to about 1800 nm, with a duration of up to about 10 seconds. For light delivered in the wavelength range of about 400 nm to about 1400 nm (visible / NIR light), the maximum flux can be about 0.6 J / cm2with a duration of up to about 10 seconds, and about 0.2 J / cm2with a duration in the range of about 10 seconds to about 30,000 seconds. For extended exposure, the light delivered is limited to a maximum power density (W / cm2) according to ANSI standards: visible / NIR light is limited to 0.2 W / cm2and far-IR light is limited to about 0.1 W / cm2. Without being limited to a particular theory, extended exposure to light delivered at UV wavelengths is generally not recommended according to ANSI standards.

[0096] In another aspect, the light flux at the excitation wavelength produced by the first light source 218 can be modulated so as to provide sufficient energy to propagate through the skin in the first region 206 of the patient 202 to the exogenous fluorescent agent without photo bleaching, and to irradiate the exogenous fluorescent agent with sufficient energy to induce detectable fluorescence at the first light detector 222 and / or the second light detector 224. In another aspect, the light flux at the emission wavelength produced by the second light source 220 can be modulated so as 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 photo bleaching, to appear as detectable light at the first light detector 222 and the second light detector 224, respectively. As non-limiting examples, the light flux produced by the light sources at 450 nm or 500 nm can be limited to 1.5 mW / cm 2 and 5 mW / cm 2 , respectively, to prevent photo bleaching.

[0097] In various aspects, the light flux produced by the first light source 218 and the second light source 220 can be modulated by any suitable system and / or device, without the limitations as described above herein. The modulation can be enabled once during the operation of the system 200, and thus, the light flux produced 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, the light modulation can be enabled at discrete times during the operation of the system 200, or the light modulation can be continuously enabled during the operation of the system 200.

[0098] In one aspect, when the system 200 is configured in the engineering mode, the light flux can be modulated via manual adjustment of any of the power supply settings and / or optical device settings as described above. In another aspect, the light flux can be automatically modulated via one or more control schemes encoded in the light source control unit of the controller 212, as described below. In this aspect, the degree of modulation can be specified based at least in part on feedback measurements obtained by various sensors disposed in the sensor head 204 of the system 200, including but not limited to the additional light detector 226 and the temperature sensor 228, as described in greater detail below.

[0099] In various aspects, the light produced by the first light source 218 and the second light source 220 is further characterized by a pulse width, defined herein as the duration of the light produced. While pulse width is typically used to characterize the performance of light sources that produce light in discrete pulses, such as pulsed lasers, it should be understood that the term "light pulse" as used herein refers to any discrete pulse of light produced by a single light source at a single wavelength, to enable acquisition of a single fluorescence measurement by the system 200. Similarly, the term "pulse width" as used herein refers to the duration of a single light pulse produced 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: the exogenous fluorescent agent or other chromophore within the tissue of the patient 202 in the absence of photobleaching, to deliver sufficient light energy to elicit detectable fluorescence from the exogenous fluorescent agent; compliance with safety standards, such as ANSI standards, for light delivered to the patient; delivery of light at a sufficiently high rate to enable acquisition of data at a rate compatible with real-time monitoring of renal function; performance capabilities of the selected light sources, light detectors, and other devices of the system 200; preservation of the useful life of the light sources, light detectors, and other devices related to the production and detection of light energy; and other relevant factors.

[0100] In various aspects, the pulse width of the light produced by the first light source 218 and the second light source 220 can be independently selected to be a duration ranging from about 0.0001 seconds to about 0.5 seconds. In various other aspects, the pulse width of the light produced by the first light source 218 and the second light source 220 can be independently selected to be a duration ranging from about 0.0001 seconds to about 0.001 seconds, from about 0.0005 seconds to about 0.005 seconds, from about 0.001 seconds to about 0.010 seconds, from about 0.005 seconds to about 0.05 seconds, from about 0.01 seconds to about 0.1 seconds, from about 0.05 seconds to about 0.15 seconds, from about 0.1 seconds to about 0.2 seconds, from about 0.15 seconds to about 0.25 seconds, from about 0.2 seconds to about 0.3 seconds, from about 0.25 seconds to about 0.35 seconds, from about 0.3 seconds to about 0.4 seconds, from about 0.35 seconds to about 0.45 seconds, and from about 0.4 seconds to about 0.5 seconds. In one aspect, the pulse width of the light produced by the first light source 218 and the second light source 220 are both about 0.1 seconds, as Figure 5 schematically illustrated in FIG. 1.

[0101] 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 a light source per second. While pulse rate is typically used to characterize the performance of light sources that generate 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 generates discrete pulses of light at a single wavelength that is associated with the acquisition of a fluorescence measurement 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 the delivery of light to a patient; the performance capabilities of the selected light sources, light detectors, and other devices of the system 200; the rate at which light is delivered is compatible with a data acquisition rate that is fast enough for real-time monitoring of kidney function; preservation of the useful life of the light sources, light detectors, and other devices related to the generation and detection of light energy; and any other relevant factors.

[0102] In various aspects, the light sources are configured to deliver light into the tissue of the patient 202 at a single location, such as the first region 206, as Figure 2 illustrated schematically. In one aspect, the delivery of light at both the excitation wavelength and the emission wavelength to the same first region 206 enables the two pulses of light 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.

[0103] In one aspect, the first light source 218 and the second light source 220 are operably 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 operably coupled to a first optical fiber and a second optical fiber, respectively, and the first optical fiber and the second optical fiber can be connected to a third optical fiber that is configured to direct light from the first optical fiber and / or the second optical fiber into the first region 206 of the patient 202. In another aspect, the first light source 218 and the second light source 220 are operably coupled to a common optical fiber or other optical component that is configured to direct light from the first light source 218 and / or the second light source 220 into the first region 206 of the patient 202. In this aspect, the light generated by the first light source 218 and the second light source 220 can be directed into the common optical fiber or other optical component using adjustable optical devices including but not limited to a dichroic mirror or a rotating mirror in an alternating pattern.

[0104] In one aspect, the system 200 can include a sensor head 204 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 can be attached in a predetermined arrangement. In one aspect, as shown in Figure 9 and Figure 10 , the first light source 218 and the second light source 220 can 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 can 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, which are operably coupled to a single light transmission aperture 1002 (see Figure 10 ) formed through the aperture plate 702 that ensures that light of both wavelengths (i.e., excitation and emission) enters the skin of the patient 202 at approximately the same location (including but not limited to the first region 206), as schematically shown in Figure 2 . 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 output power from the LED light sources, as described in further detail below.

[0105] In one aspect, only a portion of the light energy produced by the LED light sources is transmitted to the skin of the patient 202 via the single light transmission aperture 1002. In one aspect, the skin of the patient 202 receives about 1% of the light energy produced by the LED light sources. In various other aspects, the skin of the patient 202 receives about 2%, about 3%, about 4%, about 5%, about 7.5%, about 10%, about 20%, and about 50% of the light energy produced by the LED light sources. Without being limited to any particular theory, the portion of light produced by the LED light sources that is transmitted to the skin of the patient 202 can be increased by incorporating additional optical elements configured to focus and / or direct the light from each LED light source to the light transmission aperture 1002. In another aspect, a diffuser can 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.

[0106] ii) light detector

[0107] Referring again to Figure 2In 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 aspect, 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. In conjunction with the tissue of the patient 202 being illuminated with light at the excitation wavelength only and light at the emission wavelength only in an alternating fashion (see Figure 5 Measurements from the first light detector 222 and the second light detector 224 can be analyzed as described below to measure fluorescence of the exogenous fluorescent agent, and to correct the fluorescence measurements by removing the effects of diffuse reflection of light according to the correction methods described below.

[0108] 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, each separated from the first region 206 by a nominal separation distance by which light generated by the first light source 218 and the second light source 220 is delivered to the first region 206. This nominal separation distance can be selected to balance two or more effects that can impact the quality of data detected by the light detectors. Without being limited to any particular theory, as the nominal separation distance increases, the total detected signal from the light detectors can decrease due to scattering of light along the longer optical path between the light source and the light detector. This effect can be mitigated by the selection of the emission wavelength, which can result in a less pronounced decrease in the detected fluorescence signal (i.e., light at the emission wavelength) relative to the signal associated with light detected at the excitation wavelength as the nominal separation distance increases. The longer the nominal separation distance, the higher the sensitivity to changes in signal due to changes in the optical properties of the tissue.

[0109] In one aspect, the nominal separation 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 separation 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 separation distance can be 0 mm, about 1 mm, about 2 mm, about 3 mm, about 4 mm, about 5 mm, about 6 mm, about 8 mm, and about 10 mm. In one aspect, the nominal separation distance can be about 4 mm to balance these competing effects of the logarithmic decrease in signal and the decreasing size of the background signal relative to the signal from the exogenous fluorescent agent.

[0110] Referring again to Figure 9 , the first light detector 222 can be positioned within the first detection well 908 of the sensor mount 912, and the second light detector 224 can be positioned within the second detection well 910 of the sensor mount 912 within the sensor head 204. The first light detector 222 and the second light detector 224 can receive light from the tissue of the patient 202 through the first detector aperture 1004 and the second detector aperture 1006, respectively. In one aspect, the first detector aperture 1004, the second detector aperture 1006, and the light transmission aperture 1002 are separated from one another by the nominal separation distance disclosed herein above, including but not limited to a nominal separation distance of 4 mm. In one aspect, the first detection well 908, the second detection well 910, and the light source well 902 of the sensor mount 912 can be optically isolated from one another 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 fluorescent signal from the exogenous fluorescent agent can be distinguished from the unfiltered excitation light.

[0111] In one aspect, the three apertures 704 (see Figure 7 ) of the aperture plate 702 are circular, having a diameter ranging from about 0.5 mm to about 5 mm. In various other aspects, the diameter of the apertures can range from about 0.5 mm to about 1.5 mm, about 1 mm to about 2 mm, about 1.5 mm to about 2.5 mm, about 2 mm to about 3 mm, about 2.5 mm to about 3.5 mm, about 3 mm to about 4 mm, about 3.5 mm to about 4.5 mm, and about 4 mm to about 5 mm.

[0112] In one aspect, the three apertures 704 of the aperture plate 702 are circular apertures having a diameter of about 1 mm. This limited width of the apertures can result in an effective source-detector separation that is less than the nominal separation distance due to the logarithmic drop-off of the signal as the separation distance from the light source at the skin interface of the sensor head 204 increases.

[0113] In various aspects, the light detectors 222 / 224 of the system 200 can be any suitable light detection device, without limitation. Non-limiting examples of suitable light detection devices include photomissive detectors such as photomultiplier tubes, phototubes, and microchannel plate detectors; photovoltaic detectors such as reverse-biased LEDs to act as photodiodes, photoresistors, photodiodes, phototransistors; and any other suitable light detection device. In an aspect, the light detectors 222 / 224 are sufficiently sensitive to detect fluorescence emitted by an exogenous fluorescent agent within the tissue of the patient 202, including a range of about 1% to about 40% melanin in the epidermis and a range of about 0.5% to about 2% blood volume of the skin volume. In an aspect, the light detectors 222 / 224 can be silicon photomultiplier (SPM) devices.

[0114] In an aspect, the first light detector 222 can be configured to detect light at both the excitation frequency and the emission frequency, and the second light detector 224 can be configured to detect light only at the emission frequency. In one aspect, the second light detector 224 can only respond to light of the emission wavelength due to the design and materials of the sensor elements of the second light detector 224. In another aspect, the second light detector 224 can respond to a wider range of light wavelengths, but can be downstream of an optical filter configured to only pass a portion of the incident light having the emission wavelength and further configured to block the passage of light of wavelengths outside of the emission wavelength.

[0115] Any suitable optical filter can be selected for use with the second light detector 224 to selectively detect light at the emission wavelength. Non-limiting examples of suitable optical filters include absorptive filters and interference / dichroic filters. Without being limited to any particular theory, absorptive filters do not significantly vary in performance with the angle of the incident light, whereas interference / dichroic filters are sensitive to the angle of the incident light and can require additional collimating optics to effectively filter the Lambertian light distribution representative of the light emitted from the skin of the patient 202.

[0116] In one aspect, the second light detector 224 can be downstream of an absorptive long-pass filter configured to pass light greater than a predetermined wavelength to the second light detector 224. As a non-limiting example, the second light detector 224 can be downstream of a long-pass OG530 filter configured to pass light having a wavelength greater than about 530 nm. Other non-limiting examples of suitable filters include a Hoya O54 filter and a Hoya CM500 filter.

[0117] In various aspects, an absorption filter 244 configured to absorb light at the excitation wavelength can be positioned within the second detection well 910 between the second light detector 224 and the second detector aperture 1006. In one aspect, the absorption filter 244 can be constructed from OG530 Schott glass. The thickness of the absorption filter 244 can be selected to provide an optical density sufficient to filter the excitation light by approximately three orders of magnitude. In one aspect, the thickness of the absorption filter 244 can be in the range of about 1 mm to about 10 mm. In various other aspects, the thickness of the absorption filter 244 can be in the range of about 1 mm to about 8 mm, about 2 mm to about 6 mm, and about 3 mm to about 5 mm. In various additional aspects, the thickness of the 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 from OG530 Schott glass.

[0118] In another aspect, an optical diffuser can be disposed within the light source well 902. In this aspect, the optical diffuser can mix light from the first light source 218 / 220 and the second light source 218 / 220 that enter the light source well 902. By using the optical diffuser to mix the light from the first light source 218 / 220 and the second light source 218 / 220 prior to illuminating the first region 206 of the patient 202, the similarity of the optical paths taken by the emission wavelength light and the excitation wavelength light of the patient's tissue is enhanced relative to the corresponding optical paths taken by unmixed light, thereby reducing potential sources of variation.

[0119] In one aspect, a transparent material configured to pass light at both the excitation wavelength and the emission wavelength can be located within the first detection well 908 between the first light detector 222 and the first detector aperture 1004. In this regard, the transparent material can be any material having similar optical properties as the material of the absorption filter 244, including, but not limited to, thickness and refractive index. In one aspect, the transparent material within the first detection well 908 can be fused silica glass of the same thickness as the absorption filter 244.

[0120] As a non-limiting example, in Figure 3 The transmission spectrum of the OG 530 filter is provided in . Figure 3 As shown, the transmission spectrum of the OG 530 filter overlaps with the emission spectrum of the MB-102 exogenous phosphor and the emission spectrum (emission wavelength) of the green LED used as the second light source 220. In addition, the transmission spectrum of the OG 530 filter does not include the emission spectrum of the blue LED used as the first light source 218 and the absorption spectrum (excitation wavelength) of the MB-102 exogenous phosphor.

[0121] In one aspect, transparent materials such as glass 246 and optical filter 244 can be secured to flanges formed within first detection well 908 and second detection well 910, respectively. Transparent materials such as glass 246 and optical filter 244 can be secured in place using an opaque and / or light-absorbing adhesive, including but not limited to black epoxy, to ensure that all light received through first detector aperture 1004 and second detector aperture 1006 passes through optical filter 244 or glass 246 before being detected by first and second light detectors 222 / 224. In another aspect, the sides of filter 244 or glass 246 can be painted black with a light-absorbing coating, including but not limited to India ink, to ensure that light does not reach first and second light detectors 222 / 224 without passing through optical filter 244 or glass 246.

[0122] In one aspect, due to the Lambertian distribution of the angles of light exiting the patient's skin, the height of the detector wells 908 / 910, combined with the diameter of the detector apertures 1004 / 1006, can limit the fraction of light emitted from the second and third regions 208, 210 of the patient's skin that reaches the active area of ​​the light detectors 222 / 224. In one aspect, the fraction of light emitted from the second and third regions 208, 210 of the patient's skin that is received by the light detectors 222 / 224 can be in the range of about 5% to about 90%. In various other aspects, the fraction of light can be in the range of about 5% to about 15%, about 10% to about 20%, about 15% to about 25%, about 20% to about 30%, about 25% to about 35%, about 30% to about 40%, about 35% to about 45%, about 40% to about 60%, about 50% to about 70%, and about 60% to about 90%.

[0123] In one aspect, for Figure 6 and Figure 7 In the illustrated sensor head 204 having apertures 1002 / 1004 / 1006 having a diameter of 1 mm, approximately 10% of the light emitted from the patient's skin surface can reach the active area of ​​the light detectors 222 / 224 to be detected. In various aspects, the sensor head 204 can further include additional optical elements, including but not limited to lenses and / or prisms, which are configured to compensate for the Lambertian distribution of light angles in order to enhance the proportion of light emitted from the patient's skin that is directed to the active area of ​​the light detectors 222 / 224.

[0124] iii) Temperature sensor

[0125] Reference Figure 2The sensor head 204 can also include one or more additional temperature sensors 228 configured to monitor the temperature of various regions within and proximate to the sensor head 204. Non-limiting examples of suitable regions whose temperature can be monitored by the one or more additional temperature sensors 228 include: the temperature at the surface of the skin of the patient 202; the temperature proximate to the first and / or second light sources 218, 220; the ambient temperature outside of the sensor head 204; the temperature of the housing 600 of the sensor head 204; and any other suitable region. In one aspect, the additional temperature sensors 228 can be configured to monitor the temperature proximate to 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, one or more temperatures measured by the one or more additional temperature sensors 228 can be used as feedback in the control methods for one or more temperature sensitive devices of the system 200 described below.

[0126] As a non-limiting example, the temperature measurements can be used to control the amount of light energy produced by the LEDs used as the first or second light sources 218 / 220. In this example, the LED temperature measured by the second temperature sensor 1108 (see Figure 11 ) can be used in a control scheme to modulate the amount of electricity provided to the LED light source to compensate for the effect of LED temperature on LED light output. In another aspect, the additional temperature sensors 228 can monitor the temperature of the LED light sources 218 / 220 to monitor and / or compensate for temperature variations of the LEDs, as well as to monitor and / or compensate for the temperature dependent transmission of the optical filter to maintain a relatively constant output wavelength.

[0127] As another non-limiting example, the additional temperature sensors 228 can be included in the sensor head 204 in the form of a thermistor 816 (see Figure 8 ) configured to monitor the temperature of the housing 600 proximate to the contact face 606 of the sensor head 204. With reference to Figure 7 , Figure 8 and Figure 9 , in one aspect, the thermistor 816 can be epoxied into the temperature sensor opening 706 in the aperture plate 702. In this aspect, the space 918 between the circuit board (not shown) and the lower housing 604 can be filled with a thermally conductive putty to ensure good thermal conduction and dissipation.

[0128] 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, the additional temperature sensors 228 can monitor the temperature of the LED light sources 218 / 220 to monitor and / or compensate for temperature variations of the LEDs, enabling the LED light sources 218 / 220 to maintain a relatively constant output wavelength.

[0129] In another aspect, if an over-temperature condition is detected, the temperature measured by one or more of the additional temperature sensors 228 can be used to ensure subject safety by disabling one or more electrical devices including the light sources 218 / 220 and / or the light detectors 222 / 224. In one aspect, an over-temperature condition can be indicated if the housing temperature detected by the thermistor 816 is greater than about 40 °C. In various other aspects, an over-temperature condition can be detected if the housing temperature is greater than about 40.5 °C or greater than about 41.0 °C.

[0130] B. Controller

[0131] Referring again to Figure 2 In various aspects, the system 200 can include a controller 212 configured to operate the light sources 218 / 200 and the light detectors 222 / 224 in a coordinated manner to obtain a plurality of measurements for acquiring fluorescence of an exogenous fluorescent agent 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 indicative of renal function of the patient 202. Figure 11 is a schematic diagram of an electronic circuit 1100, which in one aspect illustrates an arrangement of various electrical components that can operate the system 200. In one aspect, the controller 212 can be a computing device that further includes an operations unit 214 and a display unit 216.

[0132] i) Light Source Control Unit

[0133] Referring again to Figure 2 The controller 212 can include a light source control unit 230 configured to operate the first light source 218 and the second light source 220 to produce light at the excitation wavelength and the emission wavelength, respectively, in a coordinated manner to produce a repeating sequence of pulses as Figure 5Schematically illustrated. In various aspects, the light source control unit 230 can generate a plurality of light control signals that encode one or more light control parameters, including, but not limited to: activation or deactivation of each light source; the relative timing of activation and deactivation of each light source to enable light pulse width, pulse repetition rate, electrical power delivered to the light source, or other parameters associated with light pulse flux or light pulse power; other light source-specific parameters that control the light output of the light source; and any other relevant light control parameters. In one aspect, the light source control unit 230 can receive one or more feedback measurements that are used to modulate the plurality of control signals to compensate for variations in the performance of the light source, thereby maintaining a relatively stable output of light from the light source. Non-limiting examples of feedback measurements used by the light source control unit 230 include: light output of the light sources 218 / 220 measured within the source well 902 by the first monitor photodiode 904 and the second monitor photodiode 906, respectively; the temperature of the light sources 218 / 220; and any other feedback measurements relevant to monitoring the performance of the light sources 218 / 220.

[0134] As a non-limiting example, the light source control unit 230 can be configured to operate the LED light sources 218 / 220. In this example, the light output of the LED light sources 218 / 220 can be controlled by controlling the current amplitude provided to each LED. In one aspect, the light source control unit 230 can include at least one waveform generator 1122, which includes but is not limited to a field programmable gate array FPGA having a 16-bit DAC 1124, which is operably coupled to the LED current source 1126, such as Figure 11 In one aspect, a waveform generated by at least one waveform generator 1122, including but not limited to a square wave, can control the output from the LED current source 1126. In one aspect, the current amplitude provided to the LED light sources 218 / 220 can be adjusted based on the waveform signal provided by the waveform generator / FPGA 1122.

[0135] Reference Figure 5 In one aspect, each light pulse train 500 includes an emission wavelength light pulse 502 and an excitation wavelength light pulse 504, both of which are composed of a plurality of square waves 506 generated by the first and second LED light sources 218 / 220. Figure 11 , the square wave generated by waveform generator 1122 is received by LED current source 1126. The current generated by the LED current source includes a square wave similar to the waveform generated by waveform generator 1122. Without being limited to any particular theory, since the intensity of the light generated by the LED light source 218 / 220 is proportional to the amplitude of the received current, the light generated by the LED light source 218 / 220 also includes the following: Figure 5square wave. In another aspect, as discussed in detail 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 effects of various confounding factors, including but not limited to detection of ambient light by the detector signals generated by the light detectors 222 / 224 during illumination of the patient tissue at the emission wavelength and the excitation wavelength by the first light source 218 / 220 and the second light source 218 / 220, respectively.

[0136] In various other aspects, various alternative LED pulse modulation schemes can be equivalently employed without limitation. In one aspect, the excitation pulses and the emission pulses are delivered in alternating series, with a dark period separating each pulse. In another aspect, the first and second LED light sources 218 / 220 are each modulated at a 50% duty cycle but at different modulation frequencies, allowing the signals associated with the excitation pulses and the emission pulses to be separated by frequency filtering.

[0137] Without being limited to any particular theory, the total optical power delivered to the patient's skin can be limited by at least two factors: photobleaching of exogenous fluorophores and / or endogenous chromophores, and overheating of the patient tissue illuminated 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 about 9 mW on the optical power that can be delivered to the skin. In another aspect, 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 that remains relatively constant as long as no photobleaching of the chromophores occurs. This constant autofluorescence background can be subtracted from the raw fluorescence signal, but if the autofluorescence changes over time due to photobleaching, this background correction can 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.

[0138] 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. Because the light intensity reaching these monitor photodiodes 904 / 906 is typically much stronger than the light intensity reaching the light detectors 222 / 224 through the patient's skin, less sensitive light detection devices, including but not limited to PIN photodiodes, can be used to monitor the output of the light sources 218 / 220.

[0139] In various aspects, the system 200 can be configured to operate over a range of skin tones observed in a human population. Without being bound to any particular theory, variations in skin tone between different patients 202 can result in variations in detected fluorescent signals over a range of about three orders of magnitude in amplitude. In addition, variations in concentration of exogenous fluorescent agent within each patient 202, due to elimination of the agent in the kidneys over time, can vary over a range of about two orders of magnitude in amplitude. In various aspects, the system 200 can be configured to detect fluorescence from endogenous fluorescent agents over a range of intensities that exceeds five orders of magnitude. In these various aspects, the system 200 can be configured by modulating at least one operating parameter including, but not limited to, the amplitude of the light output of the light source 218 / 220 and the sensitivity of the light detectors 222 / 224 corresponding to the detector gain.

[0140] In one aspect, the intensity of the light output by the light source 218 / 220 can be manually set by a user via the operating unit 214. In another aspect, the light source control unit 230 can be configured to automatically modulate the intensity of the light produced by the light source 218 / 220. In one aspect, the light source control unit 230 can be configured to control the intensity of the light produced by the LED light source 218 / 220 over a range of normalized output intensity from 0 (off) to 1 (maximum power). In one aspect, the intensity of the light source 218 / 220 can be set by the light source control unit 230 in coordination with the detector gain of the light detectors 222 / 224 set by the light detector control unit 232, as described below.

[0141] In one aspect, the light source control unit 230 can use the signals obtained by the system 200 during the first 10 detection cycles after data acquisition initialization, but prior to injection of exogenous fluorescent agent, to automatically adjust the light intensity produced by the LED light sources 218 / 220, as well as the gain of the light detectors 222 / 224. In this example, the initial detection cycles can be obtained with the LED light sources 218 / 220 set to approximately 10% of maximum LED intensity (corresponding to a normalized output intensity of 0.1) and the light detectors 222 / 224 set to low gain. Based on the detected intensity of the light received by the light detectors 222 / 224 at the excitation and emission wavelengths of the one detection cycle, the corresponding LED intensity can be modulated so that the analog signals produced by the light detectors 222 / 224 can correspond to approximately ¼ of the full range of each detector analog-to-digital converter (ADC) at the low detector gain setting. If the signals produced by the light detectors 222 / 224 are inconsistent in response to the light produced 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-described method results in modulation of the LED intensity setting above the maximum intensity (corresponding to a normalized output intensity of 0.1), the LED intensity setting is set to the maximum setting. Without being bound to any particular theory, the target level of the signals produced by the light detectors 222 / 224 (i.e., ¼ of the ADC range) is selected to preserve additional light detection capability to detect signals resulting from changes in the optical properties of the tissue of the patient 202 during the study due to any one or more of a number of factors, including but not limited to the introduction of exogenous fluorescent agent into the patient 202.

[0142] In one aspect described above, once the light source control unit 230 and the light detector control unit 232 have coordinated the LED intensity settings with the detector gain of the light detectors 222 / 224 over the first 10 detection cycles, an additional 10 detection cycles are obtained to confirm the suitability of these settings for operation of the system 200 given the tissue properties of the particular patient 202, after which the LED intensity settings and detector gain are recalculated as described herein. If the newly calculated LED intensity is within two times the previously determined setting, and the detector gain does not change, the previously determined settings are maintained for subsequent data acquisition cycles used to determine renal function. Otherwise, the settings are updated using the same methods described herein, and another 10 data acquisition cycles are performed to confirm the stability of the settings. This process is repeated until either the settings are determined to be acceptably stable or 10 data acquisition cycles are performed to obtain the settings, in which case the most recently determined settings are used for all subsequent data acquisition, and the user can be notified via the display unit 216 that the settings can not be optimal.

[0143] ii) light detector control unit

[0144] Referring again to Figure 2 , the controller 212 can include a light detector control unit 232 configured to operate the first light detector 222 and the second light detector 224 to enable detection of light at the emission wavelength and unfiltered light at all wavelengths, respectively. In various aspects, the light detector control unit 232 can generate a plurality of detector control signals encoding one or more detector control parameters including, but not limited to, detector gain. In various other aspects, the light detector control unit 232 can generate a plurality of light measurement signals encoding the intensity of light detected by the light detectors 222 / 224, including but not limited to raw detector signals that can be received in various aspects by an analog-to-digital converter (ADC) 1102 (see Figure 11 ) in the system 200. In another aspect, when the system 200 is configured in engineering mode, the detector gain and / or other detector control signals can be manually set by a user.

[0145] In various other aspects, the amount of light received by the light detectors 222 / 224 can vary due to any one or more of at least several factors including, but not limited to, variations in skin tone observed between individual patients 202, variations in exogenous fluorescent agent concentration within each patient 202, and any other relevant parameters. In one aspect, the gain of the first light detector 222 and the second light detector 224 can be set by a user via the operating unit 214. In another aspect, the light detector control unit 232 can be configured to automatically modulate the gain of the light detectors 222 / 224 via the bias voltage gain of a bias voltage generator 1112 (see Figure 11 ).

[0146] In one aspect, the signals obtained during the first 10 detection cycles obtained by the system 200 after data acquisition initialization, but prior to injection of the exogenous fluorescent agent, are used by the light detector control unit 232 to automatically adjust the gain of the light detectors 222 / 224, as well as the output intensity of the light sources 218 / 220. As previously described herein, the initial detection cycles can be obtained with the LED light sources 218 / 220 set to approximately 10% of maximum LED intensity (corresponding to a normalized output intensity of 0.1) and the light detectors 222 / 224 having a low gain setting, and the LED intensity can be modulated to enable the analog signals produced by the light detectors 222 / 224 to correspond to approximately ¼ of the full range of each detector analog-to-digital converter (ADC) at the low detector gain setting.

[0147] In this aspect, if the intensity of the first LED light source 218 (generating light at the excitation wavelength) is set to the maximum of the LED power range, a high detector gain for the second light detector 224 corresponding only to filtered measurements at the excitation wavelength can be considered. In various aspects, for a given light detector, a high detector gain can be 10 times higher than the corresponding low detector gain. Without being limited to any particular theory, assuming that the exogenous fluorescent agent is MB-102 introduced into the patient 202 at a dose level of approximately 4 μmol / kg of the patient's body weight, the expected peak detected fluorescence signal from the exogenous fluorescent agent during injection and renal elimination is generally expected to be approximately 10% of the magnitude of the signal received during illumination by the first light source 218 at the excitation wavelength. In one aspect, if the expected detector signal received during maximum LED intensity illumination and with the detector gain set to a high setting is still less than 10% of the range of the detector ADC, the detector gain for that measurement is increased by a factor of 10. In another aspect, the saturation state can persist for a predefined period of time, including but not limited to a 30-second period before adjustments are made to the detector gain or LED power to avoid reacting to stray signal spikes.

[0148] On the other hand, if the detected light signal from one of the photodetectors 222 / 224 exceeds a threshold percentage of the maximum ADC range, the photodetector control unit 232 can adjust the detector gain to a lower gain level to avoid signal saturation. While the highest threshold percentage of the maximum ADC range associated with signal saturation is 100%, the onset of severe detector nonlinearity occurs at a threshold percentage of approximately 40% or higher, while the onset of mild detector nonlinearity occurs when the threshold percentage exceeds approximately 15%. In various aspects, the threshold percentage of the maximum ADC range can be 40%, 35%, 30%, 25%, 20%, 18%, 17%, 16%, 15%, 14%, 13%, 12%, 11%, 10%, 9%, 8%, 7%, 6%, or 5% of the maximum ADC range. In one aspect, if the detected light signal from one of the photodetectors 222 / 224 exceeds approximately 8% of the maximum ADC range, the gain setting is adjusted. As a non-limiting example, if the detector gain is high on a signal near saturation, it is adjusted to low. If the current detector gain is set low and the corresponding detected light signal remains above a threshold percentage of the maximum ADC range, the LED output power setting of the corresponding LED light source can be reduced by a factor of 10.

[0149] In one aspect, the photodetector control unit 232 can receive one or more feedback measurements used to modulate the plurality of detector signals to compensate for variations in photodetector performance due to variations in temperature and / or light source output. Non-limiting examples of feedback measurements used by the photodetector control unit 232 include: light output of the light sources 218 / 220 as measured within the source well 902 by the first monitor photodiode 904 and the second monitor photodiode 906, respectively (see FIG. Figure 11 ), the temperature of the light detector 222 / 224 measured by the first temperature sensor 1106, the LED temperature measured by the second temperature sensor 1108, the temperature of the sensor head housing measured by the third temperature sensor 1128, the LED supply current from the LED current source 1126, and any other feedback measurements relevant to monitoring the performance of the light detector 222 / 224.

[0150] In various aspects, the light detectors 222 / 224 can be silicon photon multiplier (SPM) detectors, which can include low-noise internal amplification and can operate at lower light levels relative to other light sensor devices such as PIN photodiodes. The detector signals generated by the SPM detectors 222 / 224 can be amplified using transimpedance amplifiers 1120 / 1118, respectively (see FIG. Figure 11 ) to convert the current generated by each SPM photodetector 222 / 224 into a measurable detector voltage. The transimpedance amplifier 1118 on the second SPM photodetector 224 (i.e., detecting only filtered light at the excitation wavelength) can include a switchable detector gain that can select a low gain configured to detect a larger dynamic range for fluorescence measurement when the first LED light source 218 is activated to generate light at the emission wavelength. When the second light source 220 is not activated, the switchable detector gain can further select a high gain setting for the second SPM photodetector 224 to enhance the sensitivity of the second SPM photodetector 224 during the phase of the detection cycle when detecting light at the emission wavelength generated by exogenous fluorescent agents within the tissue of the patient 202 to ensure that the expected dark current from the second SPM photodetector 224 occupies less than 1 / 4 of the total ADC output range. In one aspect, the second transimpedance amplifier of the second SPM photodetector 224 may include a low detector gain configured to provide a transimpedance gain of approximately 4 kΩ, corresponding to approximately twice the value of the transimpedance resistor due to differential operation, and may also include a high detector gain configured to provide a transimpedance gain of approximately 40 kΩ. In another aspect, the first transimpedance amplifier of the first SPM photodetector 222 may include a fixed detector gain configured to provide a transimpedance gain of approximately 2 kΩ.

[0151] iii) acquisition unit

[0152] Referring again Figure 2 In various aspects, the controller 212 can also include an acquisition unit 234. The acquisition unit 234 can 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 produce one or more raw signals, including but not limited to, a raw fluorescence signal encoding the intensity of fluorescence detected by the second light detector 224 during illumination at the excitation wavelength, and a raw internal reflection signal corresponding to the intensity of light at the excitation wavelength detected by the first light detector 222 during illumination at the excitation wavelength, and the intensity of light at the emission wavelength detected by both light detectors 222 / 224 during illumination at the emission wavelength.

[0153] The plurality of signals received from the various sensors and devices described above are typically analog signals including but not limited to voltage and current. In various aspects, the acquisition unit 234 can be capable of sending the analog signals to one or more analog-to-digital converters (ADCs) to convert the analog signals to 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, the analog signals encoding the intensity of light detected by the first light detector 222 and the second light detector 224 can be received by a first ADC 1102.

[0154] In various aspects, at least one 24-bit sigma-delta ADC can be used to digitize the analog signals produced by the light detectors 222 / 224 and various monitor sensors. Referring again Figure 11In one aspect, a high-speed 24-bit sigma-delta ADC 1102 can be used to digitize analog signals encoding measurements from time-sensitive sensors. In this aspect, time-sensitive sensors include sensors associated with the generation and detection of light pulses that are characterized by potentially rapidly changing signals. Non-limiting examples of 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 sigma-delta ADC 1104 can be used to digitize analog signals encoding measurements from less time-sensitive sensors. In this other aspect, less time-sensitive sensors include sensors associated with monitoring system conditions that are characterized by typically slowly changing signals, including but not limited to temperatures of various system components and / or regions. Non-limiting examples of less time-sensitive sensors of system 200 include first and second thermistors 1106 / 1108 configured to monitor temperatures of light sensors 222 / 224 and light sources 218 / 220, respectively, and a third temperature sensor 1128 configured to monitor a temperature of the housing 600 of the sensor head 204.

[0155] In various aspects, the acquisition unit 234 can also be configured to enable synchronous detection of light by the detectors 222 / 224. Without being bound to any particular theory, the synchronous detection method is believed to suppress noise from the detector signal associated with detection of light produced by the light sources 118 / 120 and fluorescence produced by exogenous fluorescent agents within the tissue of the patient 202 by differentiating the detector signal from noise associated with detecting ambient light or other sources of interference.

[0156] Figure 12 is a schematic illustration 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 produced by the LED current source 1126, also characterized by a waveform that is proportional to the analog converted square wave, drives the LED light source 218 / 220. Light produced by the LED light source 218 / 220, after passing through the tissue of the patient 202, is detected by the light detectors 222 / 224 along with fluorescence produced by the endogenous fluorescent agent, and digitized by the high-speed ADC 1102.

[0157] 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 DAC 1110 (see Figure 11) into an in-phase reference sinusoidal wave 1210 and an out-of-phase / quadrature reference cosine wave 1212. In one aspect, the digitized detector signal from the ADC 1102 and the in-phase reference sinusoidal 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 samples from the reference waves 1210 / 1214 by an amount equal to the relative delay between the DAC 1124 generating the reference waves 1210 / 1214 and the ADC 1102 of the digitized detector signal to synchronize the reference waves 1210 / 1214 with the detector data being acquired.

[0158] 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 raw digitized detector signal can also be summed in a second accumulator 1220 to generate an average intensity signal 1226. Further, the in-phase intensity signal 1224 and the quadrature intensity signal 1228 can be square root summed to generate an amplitude signal 1230.

[0159] Without being limited to any particular theory, the integration interval of the accumulators 1218 / 1220 / 1222 can correspond to an integer number of modulation periods (corresponding to the period of the digital square wave 1202) to avoid biasing the measured signal. The phase accumulators 1218 / 1220 / 1222 used to control the synchronous detection are operated 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 by 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 one part in 10 6

[0160] In one aspect, the digital square wave 1202 used to modulate the LED light source 218 / 220 and implement the synchronous detection method described herein above 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 as compared to a pure sinusoidal wave that is the same peak power level modulation waveform to enable enhanced signal-to-noise ratio (SNR).

[0161] ​In another aspect, the acquisition unit 234 can also be configured 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 can pick out each component at the fundamental, characterized by an amplitude that is (4 / π) times the amplitude of the square wave 1202 used to modulate the intensity signal 1224 / 1226 / 1228. In various aspects, to suppress 50 / 60 Hz electrical noise generated by alternating current power supplies, and corresponding 100 / 120 Hz optical noise generated by ambient light sources powered from these supplies, the integration period of the accumulators 1218 / 1220 / 1222 can be selected to be 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.

[0162] iv) Processing Unit

[0163] Referring again to Figure 2 In various aspects, the controller 212 can also include a processing unit 236 configured to apply a correction to the demodulated detector signal, and transform a selected portion of the corrected detector signal into a measure of renal function. Figure 13 is a block diagram illustrating sub-units of the processing unit 236 in an aspect. Referring to Figure 13The processing unit 236 can include a pre-processing 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 changes, power supply variations provided to the light sources 218 / 220, non-linearities in the detector response, ambient temperature variations, and tissue heterogeneity. The processing unit 236 can also include a baseline subtraction subunit 1304 configured to remove a portion of the detector signals attributable to extraneous factors such as tissue autofluorescence and / or leakage of light at the excitation wavelength through the optical filter 244 of the second light detector 224. The processing unit 236 can additionally include a diffuse reflectance correction subunit 1306 configured to enable application of a diffuse reflectance correction method to remove the effects of diffuse reflection of light within the tissue of the patient 202. The processing unit 236 can also 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 can also include an RDTC calculation subunit 1310 configured to transform the detector signals obtained during the post-reagent administration period to produce a renal decay time constant indicative of the renal function of the patient. The processing unit 236 can also include a fault detection subunit 1312 configured to monitor the magnitude of the detector signals to detect any faults of the system.

[0164] a) pre-processing subunit

[0165] In one aspect, the raw signals corresponding to the light intensities detected by the light detectors 222 / 224 corresponding to the illumination by the first and second light sources 218 / 220 at the excitation and emission wavelengths, respectively, are pre-processed using various modules of the pre-processing subunit 1302 to remove the effects of a plurality of confounding factors from the raw signals to produce signals that more accurately reflect the underlying signal of interest.

[0166] As a number of non-limiting examples, the intensity of the light produced by the light sources can vary due to one or more of a number of factors including, but not limited to, fluctuations in the current supplied to the light sources and variations in the ambient temperature of the light sources. The light characterized by two or more wavelengths emitted by the same source hole of the sensor head can not share the same path to the same detector. The detectors can have a thermally dependent sensitivity and gain. Furthermore, the optical filter associated with the second light detector 224 can have temperature dependent transmission characteristics.

[0167] In one aspect, the pre-processing subunit 1302 is configured to process raw signals corresponding to the light intensities detected by the first and second light detectors 222 / 224 in order to remove one or more of the measurement errors associated with the devices and elements of the system 200 and patient-specific factors, including but not limited to the multiple factors described above. Table 2: Light detector measurements after temperature and power fluctuation correction is a block diagram illustrating modules of the pre-processing subunit 1302 in one aspect. Table 3: Light detector measurements used to obtain fluorescence measurements corrected for variable tissue optical properties is a block diagram illustrating modules of the pre-processing subunit 1302a in a second aspect.

[0168] In one aspect, as shown in Figure 2 , the pre-processing subunit 1302 1) resamples the signals using the method of the resampling module 2202 described below, 2) removes saturated detector signals using the method of the detector output saturation detection and removal module 2204 described below, 3) corrects for temperature-dependent detector gain using the method of the detector temperature correction module 2206 described below, 4) corrects for signals for instrument light directionality using the method of the light directionality correction module 2208 described below, 5) corrects for filter throughput and temperature-dependent changes in fluorescence signals using the method of the filter throughput temperature correction (emission) module 2212 described below, 6) corrects for tissue inhomogeneity using the method of the tissue inhomogeneity correction module 2216 described below, 7) corrects for filter throughput and temperature-dependent changes in excitation light and signal decomposition using the method of the filter throughput temperature correction (excitation) module and signal decomposition module 2214 described below, and 8) corrects for light power variations using the method of the fractional photon normalization module 2218 described below.

[0169] In one aspect, as shown in Figure 10 , the pre-processing subunit 1302a uses the method of the detector temperature correction module 2206a described below to calculate signal size, resamples the signals using the method of the resampling module 2202a described below, removes saturated samples using the method of the detector output saturation detection and removal module 2204a described below, corrects for temperature-dependent detector gain using the method of the detector temperature correction module 2206a described below, corrects for light power variations using the method of the fractional photon normalization module 2218a described below, corrects for excitation light leakage onto the measured fluorescence signal using the filter throughput temperature correction (excitation) module and signal decomposition module 2214a described below, and corrects for fluorescence leakage onto the measured excitation diffuse reflectance signal using the filter throughput temperature correction (emission) module 2212a described below.

[0170] - Resampling module

[0171] Referring toFigure 22A and Figure 22B The pre-processing subunit 1302 / 1302a in various aspects includes a resampling module 2202 / 2202a configured to reduce signal variations associated with physiological processes of the patient 202, including but not limited to heartbeats and respiration. Generally, an acquisition sequence is characterized by alternating illumination intervals at excitation and emission separated by an unilluminated interval (i.e., a dark interval). Although both illumination intervals (excitation / emission) are time-stamped with the same time-stamp 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 bound to any particular theory, if the separation interval associated with the acquisition sequence is on the order of the separation interval between physiological events such as heartbeats or respiration, physiological noise can be introduced into the signal. In various aspects, this physiological noise can be reduced by resampling the signals associated with the excitation and emission illumination to overlap prior to subsequent processing of the signal.

[0172] As a non-limiting example, a sample sequence can include a 100 ms dark interval, a 100 ms illumination interval at an excitation wavelength, a second 100 ms dark interval, and a 100 ms illumination interval at an emission wavelength. Each sample data packet is recorded with a single time-stamp, and each sample data packet is separated by a 400 ms interval. Due to physiological signal variations, such as physiological signal variations from heartbeats, occurring on the same time scale, the 200 ms difference between the signal acquisition associated with the excitation wavelength and the emission wavelength becomes apparent in the signal. This physiological signal noise can be reduced using the pre-processing subunit 1302 by first resampling the signals associated with the excitation and emission illumination to overlap prior to performing any additional signal processing as described below. In this non-limiting example, the signal associated with the excitation illumination can be shifted forward by 100 ms, and the signal associated with the emission illumination can be shifted backward by 100 ms, resulting in overlap of the signals.

[0173] 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 functions as a form of low-pass filter.

[0174] - Detector output saturation detection and removal module

[0175] Referring again to Global error mapping method and Figure 14AIn various aspects, the pre-processing subunit 1302 / 1302a includes a detector output saturation detection and removal module 2204 / 2204a configured to detect and remove signal values that fall outside the detection range of the light detector 222 / 224. In one aspect, the pre-processing subunit 1302 compares the detected signals to a maximum ADC signal. If any signal falls within a threshold range of the maximum ADC signal using the average or peak signal value, the detector output saturation detection and removal module 2204 identifies the value and removes the value from further processing.

[0176] - detector temperature correction module

[0177] Figure 14B and Figure 14A In various aspects, the pre-processing subunit 1302 / 1302a includes a detector temperature correction module 2206 / 2206a configured to enable temperature correction to compensate for the thermosensitivity of the light detector 222 / 224. In one aspect, the intrinsic detector gain of a silicon photomultiplier (SPM) device, commonly used as a light detector, is proportional to the difference between the device breakdown voltage and the bias voltage (referred to herein as overvoltage) applied by the bias voltage generator 1112 (see Figure 14A ). In this regard, the breakdown voltage varies with temperature in a well-characterized manner. In one aspect, the temperature correction accounts for both this internal detector gain variation and additional temperature-dependent variations in the photon detection efficiency.

[0178] In one aspect, the temperature correction can be a scaling correction applied to the detector measurements, where the scaling correction is based on the measured detector temperature. In one aspect, the measured light detector 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):

[0179]

[0180] In equation (2), the monitoring temperature T is obtained from a first temperature sensor 1106 (see Figure 14A ) configured to monitor the temperature of the sensor 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 (To) are constants specific to the particular light detector device included in the system 200. As a non-limiting example, if the light detector 222 / 224 is a silicon photomultiplier (SPM) device, V breakdownmay be 24.5 V and T0may be 21 °C. In another aspect, the coefficient C used in equation (2) can be empirically derived based on measurements obtained using a constant phantom over an ambient temperature range of about 18 °C to about 26 °C v and C T .

[0181] In another aspect, the temperature portion of the gain correction is determined by equations (3) to (5).

[0182]

[0183]

[0184]

[0185] This gain correction can be applied to each signal amplitude measured by the first and second light detectors 222 / 224 as follows:

[0186]

[0187] In one aspect, the temperature corrected measured amplitudes from each detector and monitor photodiode are calculated from the square root sum of the in-phase intensity signal 1224 (I) and the quadrature intensity signal 1228 (Q) according to equation (1):

[0188]

[0189] The signal amplitudes from the light detectors 222 / 224 calculated using equation (1) are normalized by the monitor photodiode amplitudes for each measurement set corresponding to measurements obtained during illumination at the excitation 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 set is used for this normalization. In another aspect, if two monitor photodiodes 904 / 906 are located in the same source well 902 as the two LED light sources 218 / 220 (see Figure 5 ), the average of the two monitor photodiode amplitudes from the corresponding measurement set is used for this normalization.

[0190] In one aspect, the in-phase intensity signal 1224, the quadrature intensity signal 1228, and the average intensity signal 1226 (see Figure 14A ) are further processed for the number of accumulated samples and ADC scaling such that the intensity signals 1224 / 1226 / 1228 are returned as a fraction of the entire range of the high-speed ADC 1102 (i.e., a range from a minimum of 0 to a maximum of 1). The monitor photodiodes 904 / 906 (see Figure 15Ameasurements of the first and second monitor photodiodes 904, 906 are similarly scaled to be a fraction of the full range of the low speed ADC 1104.

[0191] In one aspect, G correction Power correction can be incorporated to correct for the effects of fluctuations in the LED power supply. In this regard, the signals from the first and second monitor photodiodes 904, 906 are calibrated by measuring the light output power with a power meter as the light intensity from the light sources 218 / 220 varies. Each light source 218 / 220, C source1 and C source2 The calibration coefficients for G source1 and C source2 are calculated as the detector measured milliwatts per each recorded monitor photodiode signal value. C magnitude and C correction are used to determine the absolute light output into the tissue at each wavelength.

[0192] Referring again to Figure 15A , the detector temperature correction module 2206a corrects the signal amplitude for varying intensity of the LEDs by using the LED output signal PD ex The temperature corrected detector signals are normalized to correct for the varying intensity of the LEDs. In this case, G em is modified as follows:

[0193]

[0194] - light directionality correction module

[0195] Referring again to Figure 16 , the pre-processing subunit 1302 in this aspect includes a light directionality correction module 2208 configured to correct variations in the detected signals associated with differences in scattering and absorption of light of different wavelengths by the tissue of the patient 202 during data acquisition. In one aspect, the correction term for light directionality can be obtained by acquiring data from one or more homogeneous tissue models and measuring the light directionality using a sensor configuration without the presence of an emission filter. The measured ratio of the signal detected by the first light detector 222 (Detl) to the signal detected by the second light detector 224 (Det2) is used to determine a coefficient G ex or G em associated with the illumination of light at the excitation and emission wavelengths, 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 signal acquired in a homogeneous medium is such that the signals measured by the first detector 222 / 224 and the second detector 222 / 224 are equivalent to each other within 20%. In other aspects, the first optical detector 222 uses a coefficient G ex or G em The correction of the signal acquired in a homogeneous medium is such that the signals measured by the first detector 222 / 224 and the second detector 222 / 224 are equivalent to each other within about 10%, about 5%, about 2%, and about 1%.

[0196] - detector non-linear response correction module

[0197] Referring again Figure 15B to the pre-processing subunit 1302, in this aspect, includes a detector non-linear response correction module 2210 configured to be able to correct variations in the detection signal associated with the non-linear response of the detector. In this aspect, the amplitude data obtained by the detector 222 / 224 can be scaled using a calibration curve based on average data.

[0198] - filter throughput temperature correction (emission) module

[0199] Referring again Figure 16 to the pre-processing subunit 1302, in this aspect, includes a filter throughput temperature correction (emission) module 2212 configured to be able to correct, during the emission wavelength illumination, variations in the detection signal associated with the temperature-dependent optical characteristics of the optical filter 244 associated with the second optical detector 224. In this aspect, the signal Det2 detected by the second optical detector 224 can be corrected according to equation (8):

[0200]

[0201] In various aspects, the signal Det2 measured by the second optical detector 224 can be monitored when the ambient temperature is cycled in a range that includes the operating temperature range or a sufficiently large subset of the range to adequately determine the temperature dependence of the emission filter. These data are acquired from a homogeneous, non-phosphor mode with 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 measurements 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 aspect, the slope of Det2 / Det1 obtained over the range of ambient temperatures during emission wavelength illumination from a homogeneous, non-phosphor mode obtains the coefficient C emF,slopeT .

[0202] - tissue heterogeneity correction module

[0203] Referring again Figure 14A In this aspect, the pre-processing sub-unit 1302 comprises a tissue heterogeneity correction module 2216 configured to correct variations in the detected signal associated with the heterogeneity of the tissue interposed between the first region 206 illuminated by the light source 218 / 220 and the second and third regions 208 / 210 in which the light detectors 222 / 224 are located. In this aspect, according to equation (9), the signal Det1 corrected for the light directionality by the light directionality correction module 2208 and the signal Det2 corrected for the filter effect by the filter throughput temperature correction (emission) module 2212 are used to calculate C hetero i.e. the coefficient that corrects for the tissue heterogeneity:

[0204] C hetero = Det2 / Det1 equation (9)

[0205] - filter throughput temperature correction (excitation) and signal decomposition module

[0206] Referring again Figure 17 In this aspect, the pre-processing sub-unit 1302 comprises a filter throughput temperature correction (excitation) and signal decomposition module 2214 configured to correct variations in the detected signal associated with the temperature-dependent optical properties of the optical filter 244 associated with the second light detector 224 during excitation wavelength illumination. In this aspect, because the emission filter is configured to block light at the excitation wavelength, the filter throughput temperature correction (excitation) and signal decomposition module 2214 performs a correction for variations in the amount of excitation light leakage due to temperature-dependent changes in the optical properties of the optical filter 244. Furthermore, the filter throughput temperature correction (excitation) and signal decomposition module 2214 is able to correct the signal measured by the first light detector 222 during excitation wavelength illumination for the presence of fluorescence caused by the excitation wavelength illumination superimposed on a portion of the signal associated with the excitation wavelength illumination.

[0207] In this aspect, as shown in equation (10), the temperature-dependent variations are calculated to affect the leakage of the excitation wavelength through the optical filter 244:

[0208] C exLT = C exLT,nom + C exLT,slopeT (T-T nom ) equation (10)

[0209] In this aspect, during excitation wavelength illumination, at the nominal operating temperature T nomC is calculated from the ratio of the signals Detl and Det2 measured from the homogeneous non-fluorescing phantom during excitation wavelength illumination over the working temperature range T exLT,nom C exLT,slopeT is calculated as the slope of the signal Det2 measured from the homogeneous non-fluorescing phantom during excitation wavelength illumination over the working temperature range T.

[0210] In this regard, the filter throughput temperature correction (excitation) module and signal decomposition module 2214 also performs signal extraction to isolate the portion of the detected signal associated with diffuse reflection of excitation wavelength illumination and fluorescence. DR ex2 i.e. the amount of excitation light incident on the second light detector 224 in the absence of the optical filter 244, is not measurable due to the presence of the optical filter 244. Furthermore, the signal Detl measured by the first light detector 222 is a composite signal from both diffuse reflection and fluorescence Flrl of the excitation wavelength illumination DR ex1 . C is obtained using the tissue heterogeneity correction module 2216 as described above. Hetero The underlying signals are extracted by using the following system of equations:

[0211] Det2 = C exLT DR ex2 + Flr2 Equation (11)

[0212] Detl = DR ex1 + Flrl Equation (12)

[0213] Flr2 = C Hetero Flrl Equation (13)

[0214] DR ex2 = C Hetero DR ex1 Equation (14)

[0215] In this regard, Flr2 is determined by solving the above system of equations using only the measurable signals Detl and Det2, as follows:

[0216] Det2 = C exLT C Hetero DR ex1 + Flr2 Equation (15)

[0217] Det2 = C exLT C Hetero (Detl - Flrl) + Flr2 Equation (16)

[0218] Det2 = C exLT C Hetero Detl - C exLT C HeteroFlr1 + Flr2 Equation (17)

[0219] Det2-C exLT C Hetero Det1 = Flr2(1-C exLT ) Equation (18)

[0220]

[0221] In this regard, once Flr2 is obtained as described above, the other signals Flr1, DR ex1 and DR ex2 can be easily obtained by substituting into the above set of equations (equations (11) to (14)).

[0222] - Partial Photon Normalization Module

[0223] Referring again Figure 18 , in this regard the pre-processing subunit 1302 includes a partial photon normalization module 2218 configured to convert the detector signals into partial photon units after pre-processing as described above for use in subsequent background subtraction and intrinsic fluorescence correction algorithms as described herein. In this regard, the detector signals can be converted to photocurrent by reversing the scaling associated with the ADC and transimpedance amplifier used to acquire the detected signals to obtain the signals 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 to units of watts. The detector signals in units of watts are then proportional to the source power in units of watts measured by an additional photodetector 226 used to monitor the output of the light source 218 / 220 to obtain the number of detected partial photons.

[0224] - Optical Power Correction Module

[0225] Referring again Linear regression method and Figure 19In this aspect, the pre-processing subunit 1302 / 1302a includes a partial photon normalization module 2218 / 2218a configured to convert the detector signal into a partial photon unit after pre-processing as described above for use in subsequent background subtraction and intrinsic fluorescence correction algorithms as described herein. In this aspect, the detector signal can be converted into a photocurrent by reversing the scaling associated with the ADC and transimpedance amplifier used to acquire the detected signal to acquire 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 units of watts. The detector signal in units of watts is then proportional to the source power in units of watts measured by an additional photodetector 226 used to monitor the output of the light source 218 / 220 to obtain the number of detected partial photons.

[0226] - Excitation light leak subtraction module

[0227] Referring again to Figure 19 In this aspect, the pre-processing subunit 1302a includes a partial photon normalization module 2222 configured to perform excitation leak subtraction on the Flr meas signal. To obtain the fluorescence signal produced only by the fluorescent photons (Flr photons ), excitation leak subtraction is performed. To remove the contribution of the excitation light, the excitation leak is considered as a fraction of the diffuse reflected excitation (ExLT ) signal, where a universal calibration factor, C ExLT , determines the fraction of the signal to be subtracted from Flr meas , as follows:

[0228]

[0229] where C ExLT is a calibration factor obtained by calculating the ratio between the excitation light detected by the two detectors on the non-fluorescent optical model as described below:

[0230]

[0231] This signal is then subtracted from Flr meas to provide the fluorescence signal produced only due to the fluorescent photons as follows:

[0232] Flr photons = Flr meas - ExLT

[0233] - Fluorescence light leak subtraction module

[0234] Referring again toFigure 20 In this aspect, the pre-processing subunit 1302a includes a fluorescence crosstalk subtraction module 2224a configured to perform a fluorescence crosstalk subtraction on the Flr meas signal. To obtain diffuse reflectance, defined herein as the excitation signal due to excitation light photons (DRex photons ) only, the fluorescence crosstalk subtraction is performed. To remove fluorescence crosstalk, a calibration factor, C FlrLT is determined based on a relationship between the amount of fluorescence crosstalk observed on a database of human subject data and the tissue heterogeneity measured through a relationship between the diffuse reflectance, the emission signal

[0235]

[0236] where in one aspect pi and p2 are approximately 0.61 and 0.01, respectively, as determined by the relationship above. In another aspect, pi and p2 can assume any other values without being limited to those defined by the relationship above.

[0237] The DRex photons signal is then calculated by subtracting this portion of the measured fluorescence from the diffuse reflectance excitation signal, as follows:

[0238]

[0239] b) baseline subtraction subunit

[0240] Referring again to Figure 19 , 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, as used herein, refers to an initial time period of measurements obtained prior to the injection of exogenous fluorescent agent. During the baseline period, it can be assumed that the fluorescence signal measured by the system 200 is associated with tissue autofluorescence and / or excitation light from the LED light source 218 / 220 that leaks through the 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 with fluorescence produced by the exogenous fluorescent agent within the patient tissue only.

[0241] 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 prior to the diffuse reflectance correction described herein below, and the subtraction of the autofluorescence effect can be performed after the diffuse reflectance correction.

[0242] c) diffuse reflectance correction subunit​

[0243] Referring again to Figure 20 , the processing unit 236 also includes a diffuse reflectance correction subunit 1306. In one aspect, the diffuse reflectance correction subunit 1306 can correct the measured fluorescence data to remove the effects of changes in the optical properties (absorption and scattering) of the tissue of the patient 202 during monitoring of the renal extraction of an exogenous fluorescent agent within the patient's tissue. As discussed above, the optical properties of the tissue can change 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 system monitored area of interest associated with changes in the concentration of endogenous chromophores such as hemoglobin and melanin.

[0244] Without being limited to any particular theory, the fluorescence measurements obtained by the system 200 for determining renal function include the emission wavelength photons detected by the second (filtered) light detector 224. These emission wavelength photons are emitted by the exogenous fluorescent agent introduced into the patient's tissue in response to the illumination of excitation wavelength photons. The emission wavelength photons travel from the fluorescent source (i.e., the exogenous fluorescent agent) through the third region 210 of the patient's skin to the second (filtered) light detector 224. However, the emission wavelength photons detected by the second (filtered) light detector 224 can also include auto-fluorescence emitted by endogenous fluorophores within the patient's tissue such as keratin and collagen, as well as the leakage of excitation wavelength light through the optical filter 244 of the second light detector 224. The excitation wavelength photons that induce fluorescence of the exogenous fluorescent agent are generated by the first light source 218 and directed into the first region 206 of the patient's skin. If the optical properties (scattering and / or absorption) of the patient's skin change over the time interval (i.e., from a few hours to about 24 hours or more) during which the detector data for determining renal function is acquired, the accuracy of the fluorescence measurements can be affected, as discussed above.

[0245] During each measurement period of 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 of the light emanating from the second region of the patient's skin using the first (unfiltered) light detector 222 and the portion of the light emanating from the third region 210 of the patient's skin using the second (filtered) light detector 224. The intensity of the light detected by each combination of excitation and emission wavelength illumination through the first region 206 and detection by the unfiltered / filtered light detectors 222 / 224 contains not only information about the concentration of the exogenous fluorescent agent in the patient's tissue, but also information about the optical properties of the patient's skin.

[0246] Figure 20

[0247]

[0248] The primary measure of fluorescence is Flr meas , the intensity of the fluorescence measured at the filtered detector.

[0249] Diffuse reflectance measurement Flr meas represents the propagation of photons to the unfiltered device and consists mainly of excitation photons.

[0250] DR em and DR em,filtered represents the propagation of emitted photons only.

[0251] Referring to Table 2, in various aspects, the light intensity measured by the second (filtered) light detector 224 during illumination by the excitation wavelength light, prior to any correction for capturing the optical properties of the tissue by the exogenous fluorescent agent (Flr meas ) is the raw intensity of the light emitted. After baseline subtraction correction as previously described herein, it is assumed that the light contained in Flr meas The emission wavelength in the range is derived primarily from exogenous fluorophores, with only a minor contribution from autofluorescence of endogenous chromophores, and is therefore referred to as Flr agent In one aspect, if it is assumed that the optical properties of the patient's skin do not change, all autofluorescence contributions will be subtracted during the baseline correction described above.

[0252] However, if the optical properties of the patient's skin change during data acquisition, the patient's skin may exhibit slightly more or less autofluorescence at the emission wavelength, thereby introducing uncertainty into the accuracy of the previously performed background subtraction correction. Furthermore, the changing optical properties of the skin may further alter the intensity of light reaching the exogenous fluorescer at the excitation wavelength, thereby altering the amount of energy absorbed by the exogenous fluorescer and the intensity of the induced fluorescence emitted from the exogenous fluorescer in response to illumination with the excitation wavelength light. In various aspects, the remaining three optical measurements enable monitoring of the patient's skin's optical properties and provide data that can be used to adjust for any changes in the patient's skin's optical properties.

[0253] Referring again to Table 2, the representative signal DRex, which has been corrected for variations in temperature and optical output, before diffuse reflection correction is applied, is meas and Flr meas , is further processed into a signal attributable only to photons of the desired wavelength. The number of photons that fall on any detector due to diffuse reflection, excitation, or fluorescence depends on the light directionality and the gain of the detector at the detection wavelength as follows:

[0254] DRex meas =A1*DRex photons +B1*Flr photons

[0255] Flr meas = A2 * DRex photons + B2 * Flr photons

[0256] where the coefficients A1, A2, B1, B2 are composed of the directivity and gain factors, for example,

[0257] A1 = d450 SPM1 * G SPM1@450

[0258] The isolation of the signals generated from the fluorescence emission and diffuse reflection, excitation wavelength photons, is performed as follows:

[0259]

[0260]

[0261] Since the renal function monitor measures a rate independent of amplitude, there is no need for a constant term before the photonic signal (e.g. ) as shown below.

[0262]

[0263] Thus, the terms (or C ExLT ) and (or C FlrLT ) can be experimentally determined to isolate Flr photons and DRex photons , respectively.

[0264] The following table indicates the names of the signals used to represent each of the four measured signals in the diffuse reflectance correction unfold. Note that either of the described pre-processing paths can be followed to obtain the signals that can be used in the diffuse reflectance correction unfold.

[0265] Figure 19

[0266]

[0267] where any of the excitation wavelength signals can be used as an alternative to the diffuse reflectance correction obtained by any of the described pre-processing methods.

[0268] Referring again to Table 2, the light intensity measured by the first (unfiltered reference) light detector 222 during illumination by the excitation wavelength light, captures a measure of the diffuse reflection of the excitation wavelength light that has propagated through the patient's skin Although the first light detector 222 is configured to detect both excitation wavelength light and emission wavelength light, the intensity of the excitation wavelength light is orders of magnitude higher than the intensity of the emission wavelength light due to the low efficiency of light production via fluorescence. In various aspects, it is assumed that the proportion of light at the emission wavelength in can be ignored. In other aspects, the proportion of emission wavelength light in is estimated and subtracted. Without being bound to any particular theory, because the intensity of excitation wavelength light directed into the patient's skin is assumed to be relatively constant and losses due to absorption of exogenous fluorescent agent are negligible, and power corrections are made as previously described herein, are used as a baseline measurement to assess changes in the optical properties of the patient's skin relative to the excitation wavelength light.

[0269] The light intensity measured by the first (unfiltered reference) light detector 222 during illumination by emission wavelength light captures a measure of diffuse reflection of emission wavelength light propagating through the patient's skin (DR em ). Without being bound to any particular theory, because the exogenous fluorescent agent does not emit emission wavelength light during this phase of the data acquisition cycle due to the absence of excitation wavelength illumination, and because the intensity of emission wavelength light directed into the patient's skin is relatively constant and power corrected as previously described herein, DR em is used as a baseline measurement to assess changes in the optical properties of the patient's skin relative to the emission wavelength light.

[0270] The light intensity measured by the second (filtered) light detector 224 during illumination by emission wavelength light captures a second measure of diffuse reflection of emission wavelength light propagating through the patient's skin (DR em,filtered ). In one aspect, DR em,filtered is made of the same assumptions as DR em described above. In addition, DR em,filtered provides a means of assessing the heterogeneity of the optical properties of the tissue. Because DR em,filtered is measured by the second light detector 224, which is configured to detect light emanating from the patient's skin at the third region 210 (see Figure 20 ), the intensity of the light measured in DR em,filtered has propagated along a light path through the patient's skin that is different from the light path traveled by the light measured in DR em . Without being bound to any particular theory, because 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 (see Figure 19 ), it is assumed that DRem,filtered and DR em Any difference between the two is due to heterogeneity in the optical properties of the light passing through the skin via the two different paths.

[0271] In one aspect, intrinsic fluorescence (IF), defined herein as the measured fluorescence at the emission wavelength that is attributed solely to the emission of the exogenous fluorescer, can be calculated according to Equation (20):

[0272]

[0273] Factors IF, Flr, DR ex , DR ex and DR em,filtered As defined above in this article. As shown in equation (20), the diffuse reflection correction measurement signal DR ex , DR em and DR em,filtered Each of the factors is raised to power k ex 、k em and k em,filtered In one aspect, each measurement in Table 2 is corrected for power / temperature and background subtraction as described above (see Figure 13 and Figure 2 ).

[0274] In various aspects, k can be determined empirically ex 、k em and k em,filtered The value of k ex 、k em and k em,filtered Non-limiting examples of suitable empirical methods for finding suitable values ​​of , include a global error mapping method and a linear regression method, both of which are described in detail below.

[0275] In one aspect, once each power (k ex 、k em 、k em,filtered), the same set of indices can be reused for subsequent intrinsic fluorescence measurements. Non-limiting examples of applications of the systems and methods described herein in which a selected set of indices 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 indices can be updated by repeated use of the systems and methods described herein. In this further aspect, a new set of indices can be determined for each use of the systems and methods, and the stored set of indices can be periodically or continuously evaluated to assess whether an update of the indices is indicated. As a non-limiting example, if analysis of multiple sets of indices determines that the indices have not changed outside a threshold range during previous uses of the system, the system can be configured to perform measurements using a previous set of indices, an average / median value of all previous sets of indices, or any other estimate of an appropriate indices based on previous index values. In this non-limiting example, if analysis of multiple sets of previous indices determines that an index has varied outside of a threshold range, reselection of an index using one of the methods described below may be indicated.

[0276] Figure 13

[0277] In one aspect, the value of power used in equation (20) above is determined empirically using a global error surface method.

[0278] exist Figure 15A and Figure 15A A flowchart is shown in FIG. 1 , which illustrates the steps of the global error surface method 1400. In this aspect, the method includes, at step 1402, for each diffuse reflectance signal (DR ex , DR em , DR em,filtered )Select each power (k ex , k em , k em,filtered ) range of values. In various aspects, the range of values ​​for each power may be affected by any one or more of a variety of factors, including, but not limited to: the design of system 200, including the design of sensor head 204; the properties of the selected exogenous fluorescer, such as excitation wavelength / emission wavelength, absorption efficiency, emission efficiency, and initial dose concentration in the patient's tissue; the species of patient 202 and the corresponding concentration of endogenous chromophores; the location of sensor head 204 on patient 202; and any other relevant factors.

[0279] In one aspect, the method may include for each coefficient (k ex , k em , k em,filtered ) A wide range is selected and a wide search is performed. The error surface from the wide 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 region from the wide search, observing the wells in the error surface within the wide search, and repeating the analysis. The method can be iterated until a suitably fine resolution is obtained that accurately obtains the minimum error. In one non-limiting example, for a human patient, the selected range of the potential factor can be k ex [0, 2], k em [0, 4], and k em,filtered [-4, 0].

[0280] Refer again Figure 15A , can be at 1404, for each power k at 1402 ex 、k em 、k em,filtered The range of values ​​selected selects the step size. In one aspect, the step size for each factor can be selected based on any one or more of at least several factors, including but not limited to: the expected sensitivity of the IF value calculated by equation (20) to changes in each factor; the total number of suitable power combinations used to calculate the IF for the factor under consideration (including available computing resources, acceptable data processing times, or any other relevant factors); and any other suitable criteria for step size.

[0281] In all respects, 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 a single power k ex 、k em 、k em,filtered The step size may be constant for all values ​​of , but the step size selected for each power may be different between different powers. As a non-limiting example, k ex The chosen step size can be 0.01, and k em and k em,filtered The selected step size of may be 0.6. In various additional aspects, the step size within one or more powers may vary within a range of values ​​for each power. As a non-limiting example, k ex The selected step sizes of can be distributed nonlinearly about the mean. In this non-limiting example, k exThe vector of potential values of k 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 decreased over a sub-range of values of the power for which the IF prediction calculated by equation (20) is more sensitive to small changes at that power. Non-limiting examples of varying the step size appropriately over a range of values of a single power include: different step sizes selected by a user, random step sizes, a linear increase and / or decrease in step size, a non-linear distribution of different step sizes, such as a logarithmic distribution, an exponential distribution, or any other appropriate non-linear distribution of step sizes.

[0282] Referring again to FIG. 14 Figure 13 , the selected range of exponents at 1402, and the selected step size at step 1404 can be used to form a vector of potential values of k ex , k em , and k em,filtered at 1406. As a non-limiting example, assume a selected range of potential exponents of [0, 2] for k ex , [0, 4] for k em , and [-4, 0] for k em,filtered , and assume a constant step size of 0.5 for all powers, then the vector created at 1406 is:

[0283] k ex = [0.0 0.5 1.0 1.5 2.0] (5 values)

[0284] k em = [0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.0] (9 values)

[0285] k em,filtered = [-4.0 -3.5 -3.0 -2.5 -2.0 -1.5 -1.0 -0.5 -0.0] (9 values)

[0286] Referring again to FIG. 14 Figure 2 , for each combination of exponents in the vectors formed at 1406, the IF is calculated at 1408 using equation (20) from the measured values Flr, DR ex , DR em , and DR em,filtered . For each combination of exponents, a plurality of IF values is calculated at 1408, where each IF value corresponds to one of the data acquisition periods (i.e., as shown in FIG. 14, 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 would be calculated. Figure 2

[0287] ​In one aspect, a combination of multiple potential indices can be evaluated to select a combination of indices from among the multiple to assign for diffuse reflectance correction for subsequent use in calculating Equation (20). Referring again to Figure 2 An error estimate of the corrected Flr signal data (i.e., IF signal data calculated using Equation (20)) can be calculated at 1410. Any estimate of error can be calculated at 1410, including but not limited to a curve fit to the IF signal data, a quantity 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 a single exponential curve fit. Without being limited to any particular theory, it is believed that the rate of clearance of an exogenous fluorescent agent (such as MB-102) from the kidneys is expected to be a constant exponential decay characterized by a renal decay time constant RDTC.

[0288] In one aspect, a subset of the Flr signal corresponding to a post-reagent administration period 1508 / 1510 can be selected for estimating the error for each combination of indices used to calculate the IF signal using Equation (20) relative to a reference curve, including but not limited to a curve obtained using plasma measurements. As a non-limiting example, if an exogenous fluorescent agent is introduced into the tissue of a patient by way of an intravenous injection, then the post-reagent administration period includes a time period after the injection in which the exogenous fluorescent agent has undergone sufficient diffusion from the blood to the entire extracellular fluid space of the patient such that the decay of fluorescence is indicative of clearance of the agent by the kidneys. In various aspects, the post-reagent administration period 1508 / 1510 of the Flr measurements can be selected without limitation by any suitable method. Non-limiting examples of suitable methods for identifying the post-reagent administration period include selection via inspection by a user and automated selection methods, such as a plateau detection method enabled by the plateau selection subunit 1308, as described in detail below.

[0289] Figure 2 is a plot of fluorescence measurements obtained from a patient over a period of about 10 hours after injection of an exogenous fluorescent agent (MB-102) about 3 hours prior to injection. Referring to Exogenous label , the pre-injection / baseline period 1502 is characterized by a relatively low and stable level of fluorescence, likely due to the absence of endogenous fluorescent agent in the patient's blood. After injection of the exogenous fluorescent agent 1503, the fluorescence measurements show a sharp increase 1504 to a peak concentration 1506, followed by a relatively smooth exponential decrease 1508 back to a background level of fluorescence as the kidneys eliminate the exogenous fluorescent agent from the patient's blood. Without being limited to any particular theory, it is believed that the injected exogenous fluorescent agent is likely to be in equilibrium on the extracellular space once the fluorescence decay is well described by a linear fit (or a line on a semi-log plot). Examples is an enlargement of the plot of Example 1 : Perturbation analysis ,Figure 21A The measured fluorescence data is shown in comparison to a linear curve fit 1604 and a log of the IF signal over a portion of the post-balance period 1510, showing a close fit of the single exponential curve fit to the IF signal data.

[0290] In one aspect, the log of the computed IF signal values can be fit to a line, and for each of a plurality of combinations of exponents, the error of the curve fit relative to the individual IF values can be compared to the IF signal computed using equation (20) to compute an error at 1410. Any statistical summary parameter suitable for quantifying the error of the single 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 computed at 1410 can be the normalized RMS error of the linear fit of the log (IF) signal. In this aspect, the normalized RMS error computed at 1410 is a single numerical quantity to facilitate the subsequent selection of a single combination of exponents from the plurality of combinations identified at 1406.

[0291] Referring again to Figure 21A , the method 1400 includes selecting a single combination of exponents from the plurality of combinations for which IF is computed at 1408 at 1412. Without being limited to any particular theory, it is postulated that the combination of exponents associated with the computed IF signal that minimizes the error computed at 1410 is best suited to correct the measured Flr signal to eliminate the effects of changes in the optical properties of the patient's skin during data acquisition in the post-reagent administration 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 map of all error values corresponding to all combinations of exponents.

[0292] In various aspects, the plurality of error values corresponding to the plurality of combinations of exponents can be converted into an error map comprising a three-dimensional volume, where each of the three dimensions respectively correspond to the powers used in equation (20): k ex , k em , and k em,filtered . In these various aspects, each error value corresponding to one of the combinations of exponents is mapped to a coordinate (k ex1 , k em1 , and k em,filtered1 ) within the three-dimensional volume, where k ex1 , k em1 , and k em,filtered1a value that is a combination of the exponents. In various aspects, each error value can be mapped to a three-dimensional volume in any known form, including but not limited to: numbers, colors, grayscale values, and any other suitable form.

[0293] In one aspect, the three-dimensional mapping of error values described above can be transformed into a plurality of error surfaces corresponding to the plane mapping of error values associated with a single value of one of the powers k ex , k em , and k em,filtered , with the entire value range of the remaining two exponents used as the horizontal and vertical axes of the error mapping.

[0294] Figure 21B is an error mapping of the normalized RMS error of the single exponential curve fit of the calculated IF signal mapped with a constant value of k em to the entire range of k em,filtered (horizontal axis) and k ex (vertical axis), where the normalized RMS error is represented as a color on the error mapping. In one aspect, the normalized RMS value calculated for each coefficient can be normalized according to equation (21):

[0295]

[0296] where IF agent is the calculated IF signal, and fit(IF agent ) is the respective value of the single coefficient curve fit equation. In one aspect, the global error mapping method determines the powers used in the correction for changes in skin optical properties by analyzing a single measurement set as described above. In another aspect, the global error mapping method can analyze and combine multiple measurement data sets from 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 from multiple individuals obtained using multiple systems and sensor heads. In one aspect, the powers used in the correction according to equation (20) can be determined for each measurement of each individual. In other aspects, the powers to be used can be obtained using at least several different measurement data sets, and the powers so 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 can be examined, k em,filtered -k ex projections, k em,filtered -k em projections, and k em -k exprojection) to determine whether the power range defined at 1402 is sufficient. In one aspect, the error surface can be examined to confirm that the map includes a well-defined minimum. In this aspect, if the examination of the error map does not identify a minimum, the range of one or more power values ​​can be modified and the method 1400 can be repeated. In one aspect, an assessment of the optical properties of the patient's skin (e.g., melanin absorptivity, blood content, and / or scattering coefficient) can be used to classify the patient so that an appropriate coefficient set can be selected for that class.

[0297] In one aspect, the index k may be stored ex 、k em and k em,filtered and used in subsequent measurements performed by system 200. Table 4. Effect of diffuse reflectance data correction is a graph comparing the raw fluorescence signal (blue line) to the calculated IF signal (red line) for the measurement data set. In another aspect, a global correction can be calculated by combining measurements obtained using multiple different systems and / or sensor heads and identifying the combination of indices that corresponds to the overall minimum error value.

[0298] Perturbation

[0299] In one aspect, the value of the power coefficient used in equation (20) above is determined empirically using a linear regression method. Correction effect A flowchart showing the various steps of a linear regression method for obtaining a linear regression model with predictor variables (DR ex , DR em , DR em,filtered ) in the form of a regression equation.

[0300] Reference Applied pressure , method 1900 may include converting the Flr logarithm to log(Flr) to prepare the raw fluorescence measurement Flr for analysis at 1902. Data shift / outlier reduction is a graph of log(Flr) generated at 1902. Sensor lateral movement In one aspect, method 1900 may further include selecting a region of stable optical properties 2002 (see Data shift / outlier reduction ), in this respect, the region of stable optical properties 2002 generally corresponds to IF signal slope reduction improvement In this regard, method 1900 further includes obtaining a linear regression model 2004 within the region of stable optical properties 2002 at 1906. Linear regression model 2004 can be obtained using any regression method without limitation, including but not limited to multivariate linear regression modeling methods.

[0301] Refer againIF signal slope reduction improvement The method 1900 can further include, at 1908, extending the linear regression model 2004 obtained within the region of stable optical properties 2002 to produce an extended linear regression 2008 that extends into the region of variable optical properties 2010. In one aspect, the region of variable optical properties 2010 is characterized by a non-linear distribution within the log(Flr) plot shown in Sensor head removal / replacement

[0302] Referring again to Data shift / outlier reduction The method 1900 can further include obtaining a linear regression model 2004 having the predictor variable Flr, DR ex , DR em , and DR em,filtered and the linear curve fit 2004 as the predicted response. The extension of the linear regression 2008 produced at 1908 can be used to train the linear regression model obtained at 1910.

[0303] In one aspect, the linear regression model can be developed using a single set of measurement data obtained from a single individual and / or a single system and sensor head. In another aspect, the linear regression model can be developed using a plurality of sets of measurement data obtained from a plurality of individuals and / or a plurality of systems and sensor heads. In some aspects, the linear regression model can be redeveloped for each new set of measurement data obtained for an individual. In at least some other aspects, the constants and parameters characterizing the linear regression model developed as described above can be stored for subsequent redevelopment of the linear regression model in place of each set of measurement data obtained as described above.

[0304] d) fault detection subunit

[0305] Referring again to Cooling The processing unit 236 of the controller 212 can further include a fault detection subunit 1312 configured to monitor the functionality of the light source 218 / 220 and the light detector 222 / 224 and notify the user of any irregularities of any detected faults within the system 200 via the display unit 216. In various aspects, the fault detection subunit 1312 can check the signal levels received from the light source 218 / 220 and the light detector 222 / 224 and the additional temperature sensor 228 and additional light detector 226 of the associated sensor head 204 (see No significant impact ​) to achieve basic identification of fault and notification conditions. In various aspects, the signal amplitude (see Equation (1)) and the average signal can be used to determine the peak and trough levels of the modulation of the LED light source 218 / 220. The trough of the signal, defined herein as the average signal minus one-half of the peak-to-peak signal, can be used in one aspect to monitor the ambient light level. Without being bound to any particular theory, additional contributions to the trough level of the modulated signal, such as amplifier DC offset, can be negligible as a constant contribution that is small relative to the ambient light leakage. In one aspect, if the ambient light level detected at low detector amplifier gain is recorded to exceed about 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.

[0306] In various other aspects, saturation of the photodetector 222 / 224 detector can also be monitored by the fault detection subunit 1312. In these other aspects, saturation can be monitored by calculating the peak of the signal, defined herein as the average signal value plus one-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 subunit 1312 can issue a saturation notification to the user via the display unit 216. If the fault detection subunit 1312 detects a saturation event, the ambient light level can then be checked to determine if the saturation event is associated with ambient light saturation, defined herein as a saturation event that occurs concurrently with an ambient light notification as described above. If an ambient light saturation event is detected, the fault detection subunit 1312 issues an ambient light saturation notification to the user via the display unit 216, and data acquisition by the acquisition unit 234 continues in this notification state to allow the user to resolve the condition. If a saturation event is detected that is not associated with excess ambient light, the fault detection unit can signal the photodetector control unit 232 to perform adjustment of the detector gain and / or can signal the light source control unit 230 to perform 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 associated with excess ambient light. In some aspects, if a saturation event is detected, but the user has disabled automatic gain control when the system 200 is configured in engineering mode as described above, the user is also notified via the display unit.

[0307] e) Post-Reagent Application Selection Subunit

[0308] Referring again to Figure 21C , the processing unit 236 can also include a post-reagent application selection subunit 1308 configured to automatically identify portions of the measurement data set corresponding to the post-reagent application time periods 1508 / 1510 (see Figure 21C ) Referring again to Example 2: Sensor head with flared housingAs described above, after injection of an exogenous fluorescent agent, such as MB-102, into the bloodstream of a patient, the exogenous fluorescent agent undergoes a period of equilibration as it diffuses from the bloodstream into the remaining extracellular tissue of the patient. The temporal profile of the fluorescent signal, IFr, after injection of the agent 1503 can be characterized as a bi-exponential signal profile described by equation (22):

[0309]

[0310] where Co is the baseline signal, typically removed by baseline subtraction as described above.

[0311] Referring again to Figure 23 Once the exogenous fluorescent agent has diffused into the extracellular tissue of the patient to reach quasi-steady state conditions, a post-equilibration period 1510 is reached, and the fluorescent signal can be characterized as a linear decay. Without being bound to any particular theory, it is hypothesized that the post-equilibration region of the measured data set is characterized as a 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):

[0312] IF post-equilibration = Co + C1e -t / τ Equation (23)

[0313] In one aspect, the post-agent administration selector unit 1308 can automatically identify the post-agent administration period 1510 by performing single exponential curve fitting at different portions of the IF data set and analyzing the curve fitting error associated with each of the different portions. In various aspects, the post-agent administration selector unit 1308 can select the earliest occurring portion of the IF data set in which the curve fitting error associated with the single exponential curve fit is below a threshold value as the initial post-agent administration portion of the IF data set suitable for data correction and analysis as described above. Any analysis method suitable for comparing the curve fitting error associated with single exponential curve fits of different portions of the IF data set can be used in the post-agent administration selector unit 1308, including but not limited to linear curve fitting portions of the IF data set falling within overlapping or non-overlapping data windows and comparing the curve fitting error of the respective data windows. In one aspect, the post-agent administration selector unit 1308 can generate at least one signal configured to signal to the diffuse reflectance correction unit 1306 and / or the RDTC calculation unit 1310 a time range within the IF data set corresponding to the post-agent administration period 1508 / 1510 to enable selection of the appropriate portion of the IF data set for correction and analysis as disclosed herein.

[0314] In another aspect, a linear fit and a 2-exponential fit to the IF data can be compared. In this other aspect, once the fit errors are equivalent (corrected for the additional degree of freedom in the 2-exponential fit), the equilibrium can be identified as being complete.

[0315] f) RDTC calculation subunit

[0316] In various aspects, the system 200 is configured to convert, in response to illumination of light at the excitation wavelength, various measurements from the light detectors 222 / 224 and associated light sources 218 / 220 and other thermal and light sensors, into a corrected intrinsic fluorescence (IF) signal corresponding to detected fluorescence that can be attributed solely to fluorescence emitted by the exogenous fluorescent agent at the emission wavelength. In various aspects, an exponential decline of the IF signal during the post-reagent administration portion of the IF data set can be analyzed to monitor and quantify kidney function.

[0317] In one aspect, an exponential decline of the IF signal during the post-reagent administration portion of the IF data set can be converted into a glomerular filtration rate (GFR) that is configured to quantify kidney function. In another aspect, an exponential decline of the IF signal during the post-equilibrium portion of the IF data set can be converted into a renal decay time constant (RDTC) that is also configured to quantify kidney function. In another aspect, an exponential decline of the IF signal during the post-equilibrium portion of the IF data set can be converted into a renal decay rate that is also configured to quantify kidney function.

[0318] Referring again to Figure 24 , the processing unit 236 can also include an RDTC calculation subunit 1310 that 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 the 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 log-transformed IF signal data (log(IF)), as described in equation (24):

[0319]

[0320] In various aspects, the RDTC calculation subunit 1310 can generate a signal 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 plot of the RDTC as a function of time, a single discrete RDTC value, a table of RDTC values as a function of time, a color-coded display, or other graphical representation configured to determine whether the calculated RDTC can be classified as normal / healthy, abnormal, high, low, and any other suitable classification. In various other aspects, any of the aforementioned graphical forms can be updated continuously or non-continuously as additional data is obtained and analyzed. 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.

[0321] In another aspect, the RDTC calculation subunit 1310 can convert the RDTC to a glomerular filtration rate (GFR) using known methods. In this aspect, the RDTC can be inverted and multiplied by a slope, resulting in a cGFR, a prediction of GFR, which can be corrected for body size (e.g., body surface area or distribution volume).

[0322] v) memory

[0323] Referring again Figure 24 The controller 212 of the system 200 can also include a memory 242 configured to facilitate storage of data in the system 200. In some embodiments, the memory 242 includes a plurality of storage components such as, but not limited to, a hard disk drive, a flash memory, a random access memory, and a magnetic or optical disk. Alternatively or additionally, the memory 242 can include a server such as a server in communication with the controller 212. The memory 242 stores at least one computer program that, when received by the 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 tape device, a flash memory or other similar solid state memory device, or an array of devices, including devices in a storage area network or other configurations. The 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 carriers can be non-transitory computer- or machine- readable media, such as the memory 242 or memories on the processor 238.

[0324] In various aspects, the system 200 can record raw and processed data to a series of files. Each file can contain a header file containing information about the operator, tool, and session. Each experimental session records a set of files to a separate folder for each sensor head used in that session. The raw data file can contain in-phase, quadrature, and average measurements from the detector and monitor, as well as the gain settings for the LEDs and detector during the active period of the excitation and emission wavelengths LEDs.

[0325] In various other aspects, the processed data file can contain fluorescence and diffuse reflectance measurements after amplitude calculation and correction of the monitor readings, as well as the gain settings for the LEDs and detector. The intrinsic fluorescence data file can contain intrinsic fluorescence measurements resulting from diffuse reflectance correction of the raw fluorescence signal. The GFR file can contain the calculated GFR as a function of time, classified to indicate whether post equilibration occurred, as well as confidence limits. The telemetry file can contain temperature and voltage measurements. The event log file can contain user and automatically generated event logs.

[0326] vi) GUI unit

[0327] Referring again Figure 24 In various aspects, the controller 212 can include a GUI unit 240 configured to receive a plurality of signals from other units of the system, the plurality of signals encoding various measured and transformed data. Further, the GUI unit can be configured to generate signals configured to operate the display unit 216 to display data, frames, tables, and / or any other communication of information between a user and the system 200.

[0328] vii) Processor

[0329] Referring again Figure 25 The controller 212 can also include a processor 238. The processor 238 can include any type of conventional processor, microprocessor, or processing logic that interprets and executes instructions. The processor 238 can be configured to process instructions for execution within the controller 212, including instructions stored in the memory 242 to display graphical information for a GUI on an external input / output device, such as the display unit 216 coupled to the high speed interface. In other implementations, multiple processors and / or multiple buses can be used as appropriate, as well as multiple memories and types of memory. Further, multiple controllers 212 can be connected with each device providing a portion of the necessary operations to enable the functionality of the system 200. In some embodiments, the processor 238 can include the acquisition unit 234, the light detector control unit 232, the light source control unit 230, and / or the processing unit 236.

[0330] As used herein, a processor, such as processor 238, can include any programmable system including systems using microcontrollers, reduced instruction set circuits (RISC), application specific integrated circuits (ASICs), logic circuitry, and any other circuit or processor capable of executing the functions described herein. The above examples are exemplary only, and are thus not intended to limit in any way the definition and / or meaning of the term “processor.”

[0331] As described herein, computing devices and computer systems include processors and memory. However, any processor in a computer device referred to herein can also refer to one or more processors, where the processors can be in one computing device or in multiple computing devices operating in parallel. Additionally, any memory in a computer device referred to herein can also refer to one or more memories, where the memories can be in one computing device or in multiple computing devices operating in parallel.

[0332] C. Operating Unit

[0333] Operating unit 214 can be configured to enable a user to interface with controller 212 (e.g., visually, audially, touch, button press, stylus tap, etc.) to control the operation of system 200. In some embodiments, operating unit 214 can further be coupled to each sensor head 204 to control the operation of each sensor head 204.

[0334] D. Display Unit

[0335] Referring again to Figure 26 , system 200 can also include a display unit 216 configured to enable a user to view data and control information for system 200. Display unit 216 can also be coupled to other components of system 200, such as sensor heads 204. Display unit 216 can 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, display unit 216 can be configured to present a graphical user interface (e.g., a web browser and / or a client application) to a user. The graphical user interface can include, for example, a display of GFR values generated by system 200 as described above, as well as operational data for system 200.

[0336] Figure 25

[0337] Without being bound to any particular theory, highly hydrophilic and small (creatinine, molecular weight = 113) to moderately 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 by the renal tubules nor secreted, has negligible binding to plasma proteins, and has very low toxicity. To design an optical probe that meets all of these requirements, a balance is struck between the photophysical properties and the molecular size and hydrophilicity of the fluorophore. For example, while hydrophobic cyanine and indocyanine dyes optimally absorb and emit in the near-infrared (NIR) biological window (700 to 900 nm), the hydrophilicity is insufficient for use as a pure GFR agent. Smaller dye molecules can be more easily transformed into the highly hydrophilic species required for renal clearance, but the limited π-systems formed by these lower molecular weight compounds are generally capable of exciting and emitting single photons in the ultraviolet (UV).

[0338] To address the pharmacokinetic issues associated with enhancing the photophysical properties, simple derivatives of 2,5-diaminopyrazine-3,6-dicarboxylic acid have been used as a fluorescent scaffold system of very low molecular weight with bright emission in the yellow to red region of the electromagnetic spectrum. To simultaneously optimize the pharmacokinetics of GFR and the photophysical properties, a SAR study has been performed using amide-linked variants of these derivatives. A variety of hydrophilic functionalities can be employed to achieve rapid renal clearance of this class of pyrazine fluorophores, including carbohydrates, alcohols, amino acids, and various PEG-based related strategies. PEG surrogates can be used to increase hydrophilicity and solubility, reduce toxicity, and modulate the aggregation of the resulting pyrazine derivatives. A range of medium-sized PEG pyrazine derivatives can also be adapted for use as endogenous fluorescent agents, with variations in the molecular weight and structure (and hence the hydrodynamic volume) of the series.

[0339] In one aspect, the exogenous fluorescent agent is MB-102.

[0340] Figure 25

[0341] The following examples illustrate various aspects of the disclosed systems and methods.

[0342] Figure 26

[0343] To demonstrate the effectiveness of the diffuse reflectance data correction method described above herein, the following experiment was performed.

[0344] Using the methods described above herein, particularly the diffuse reflectance data correction method, a system similar to the system 200 described above herein was used to monitor the fluorescence produced during renal elimination of the exogenous fluorescent agent MB-102.

[0345] Figure 26is a plot summarizing the change in amplitude of the raw fluorescence signal (Flr) before and about 6 hours after injection of the MB-102 fluorescent agent into the pig. During the post-equilibrium portion, corresponding to ​ At a time of about 13:45 in the middle, the pig was subjected to a selected series of perturbations to change 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, Sp02 reduction 2108, Sp02 reduction 2110, removal / replacement of the sensor head 2112 / 2114, and skin cooling 2116.

[0346] ​ is a plot summarizing the corrected intrinsic fluorescence signal (IF) corrected as described above without baseline subtraction. The time course of the IF signal after the time point of 2 hours after injection of the reagent is characterized by an expected signal single exponential decay with decay changes due to the applied perturbations.

[0347] Table 4 summarizes the specific effects of the diffuse reflectance data correction method on the Flr data associated with each individual perturbation:

[0348] ​

[0349] ​ ]]> ​ ​ ]]> ​ ​ ​ ​ ​ SpO2 decrease ​ SpO2 increase ​ ​ ​ ​ ​

[0350] ​ is a plot summarizing the detected diffuse reflectance signals DR em,filtered , DR em , and DR ex that are plugged into equation (20) to determine the diffuse reflectance correction of the raw Flr signal as described above. As shown in ​ , the DR em,filtered signal is most sensitive to the various perturbations. The DR em and DR ex signals exhibit moderate changes in response to the perturbations.

[0351] The results of these experiments show that the diffuse reflectance data correction is able to correct the raw fluorescence signal data to compensate for the effects of the various perturbations that cause various changes in the optical properties of the skin.

[0352] ​

[0353] ​is 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.

[0354] ​ is a bottom view of the sensor head 204a showing the bottom surface 608a of the housing 600a. The bottom surface 608a can include an aperture plate 702a that includes one or more apertures 704a configured to transmit light between the patient's skin and the light sources and light detectors contained within the housing 600. As shown, the apertures 704a include light transmission apertures 1002a configured to transmit the illumination generated by the first and second light sources 218 / 220 to the tissue of the patient 202, and first and second detector apertures 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 apertures 704a under a relatively large area that is shielded by the bottom surface 608a from ambient light conditions. This reduction in scattered ambient light entering the first and second detector apertures 1004 / 1006 reduces the noise introduced into the light intensity measurements taken by the first and second light detectors 222 / 224. ​

[0355] In various aspects, the bottom surface 608a of the housing 600a can be attached to the patient's skin using a biocompatible and transparent adhesive material 610a including, but not limited to, a transparent double-sided medical grade adhesive, such as 3M™ Scotch-Weld™ Optically Clear Adhesive 8141, as shown. ​

[0356] ​ is an isometric view of the sensor head 204a with the upper housing 602a and various electrical components removed to reveal the inner housing 2502. ​ is an exploded view of the inner housing 2502 and associated electrical components, as shown. ​ ​ ​ ​​​​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, respectively, of the sensor mount 912.

[0357] Referring now to ​ 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, respectively, of the sensor mount 912.

[0358] 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 optically transparent conduits between the tissue and the interior of the sensor head 204a. In addition, diffusers 2616, 2618, and 2620 are coupled over the optically transparent windows 2610, 2612, and 2614, respectively. The diffusers 2616, 2618, and 2620 are configured to spatially homogenize the light delivered to the tissue by the light sources 218 / 220 and to spatially homogenize the light detected by the light detectors 222 / 224. In one aspect, the 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.

[0359] In view of the above, it will be seen that the several advantages of the disclosure are achieved and other advantageous results obtained. As various changes could be made in the above methods and systems without departing from the scope of the disclosure, it is intended that all matter contained in the above description and shown in the accompanying drawings shall be interpreted as illustrative and not in a limiting sense.

[0360] When introducing elements of the disclosure or the various versions, embodiments, or aspects thereof, the articles "a," "an," "the," and "said" are intended to mean that there are one or more elements. The terms "comprising," "including," and "having" are intended to be inclusive and mean that there can be additional elements other than the listed elements.

Claims

1. A method for monitoring a time-varying fluorescence signal emitted from a fluorescent agent within a diffusely reflecting medium having time-varying optical properties, the method comprising: A measurement data set is provided, comprising a plurality of measurement data entries, each measurement data entry comprising at least two measurement values ​​obtained jointly from the patient at one data acquisition time before and after administration of the fluorescent agent, the at least two measurement values ​​being selected from: During illumination of the diffusely reflective medium by light of an excitation wavelength from a first region adjacent to the diffusely reflective medium, the DR detected by the unfiltered light detector at a second region adjacent to the diffusely reflective medium is ex signal, wherein the DR ex a signal representing the combined intensity of all light received by the unfiltered light detector from the second region; an Flr signal detected by a filtered light detector at a third region adjacent to the diffusely reflective medium during illumination of the diffusely reflective medium by the excitation wavelength light from the first region, wherein the Flr signal represents an intensity of the emission wavelength light received by the filtered light detector from the third region; During the period when the diffuse reflective medium is illuminated by the emission wavelength light from the first region, the DR detected by the unfiltered light detector at the second region is em signal, wherein the DR em a signal representing the combined intensity of all light received by the unfiltered light detector from the second region; and The DR detected by the filtered light detector at the third region during illumination of the diffuse reflective medium by the emission wavelength light from the first region is em,filtered signal; wherein the DR em,filtered a signal representing the combined intensity of the emission wavelengths of light received by the filtered light detector from the third region; identifying a post-agent administration portion of the measurement data set; and converting each Flr signal of each measurement data entry within the post-reagent application portion of the measurement data set into an IF signal representing detected fluorescence intensity emitted solely by the fluorescer from within the diffusely reflecting medium, wherein converting comprises combining the at least two measurements according to a conversion relationship comprising a mathematical equation for converting Flr to IF; wherein at least one of the at least two measured values ​​comprises the Flr signal, wherein two light pulses at an excitation wavelength and an emission wavelength are transmitted into the patient's tissue at the first region, thereby enabling the two light pulses to share at least a portion of an optical path through the patient's tissue between an entry point at the first region and detection points at the second region and the third region.

2. The method according to claim 1, wherein The conversion relationship is composed of equation (20): Wherein, IF represents the fluorescence intensity emitted by the fluorescent agent alone, and k ex 、k em and k em,filtered DR ex , DR em and DR em,filtered Index of Among them, DR ex , DR em and DR em,filtered Indicates the diffuse reflection corrected measurement signal.

3. The method according to claim 2, wherein: k ex 、k em and k em,filtered is predetermined from previous analysis of previous measurement datasets.

4. The method according to claim 2, wherein: Determine k by global error mapping method ex 、k em and k em,filtered ,include: Form three vectors of suggested index values, each of which includes k ex 、k em and k em,filtered Multiple suggested values ​​for transforming each Flr signal for each measurement data entry within the post-reagent administration portion of the measurement data according to the transformation relationship using each combination of suggested values ​​from the three vectors to form a plurality of transformed data measurement sets; performing a single exponential curve fit on at least a portion of the measurement data entries of each of the plurality of transformed data measurement sets to obtain a plurality of curve fit errors, each curve fit error corresponding to a combination of suggested exponential values ​​from the three vectors; compiling an error map comprising at least a portion of the plurality of curve fit errors mapped to a volume defined by two or more orthogonal axes, each orthogonal axis comprising a range of suggested index values ​​from one of the three vectors; identifying a minimum curve fit error within the error map; and The suggested exponent value corresponding to the minimum curve fit error is selected for use in equation (20).

5. The method according to claim 4, wherein Each curve fit error in the plurality of curve fit errors comprises a normalized root mean square fit error of the single exponential curve fit.

6. The method according to claim 1, wherein The transformation relationship is composed of a linear regression model, which uses the low variability part of the plurality of measurement data items to be predicted by the variable DR ex , DR em and DR em,filtered forming, a low variability portion of the plurality of measurement data entries characterized by a curve fit error below a threshold for a single exponential curve fit of the measurement data entries of the low variability portion; Among them, DR ex , DR em and DR em,filtered Indicates the diffuse reflection corrected measurement signal.

7. The method according to claim 6, wherein: The linear regression model is extrapolated to measurement data entries outside the region of low variability.

8. The method according to claim 1, wherein Also included is subtracting a baseline value of Flr from each Flr value of a plurality of measurement data sets before converting each Flr signal of each measurement data entry within the post-reagent administration portion of the measurement data set to an IF signal.

9. The method according to claim 2, wherein: k ex 、k em and k em,filtered Any one or more of is equal to 0.

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