NZ755760B2 - Method for non-invasive monitoring of fluorescent tracer agent with background separation corrections - Google Patents
Method for non-invasive monitoring of fluorescent tracer agent with background separation corrections Download PDFInfo
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- NZ755760B2 NZ755760B2 NZ755760A NZ75576018A NZ755760B2 NZ 755760 B2 NZ755760 B2 NZ 755760B2 NZ 755760 A NZ755760 A NZ 755760A NZ 75576018 A NZ75576018 A NZ 75576018A NZ 755760 B2 NZ755760 B2 NZ 755760B2
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- G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
- G01J1/00—Photometry, e.g. photographic exposure meter
- G01J1/42—Photometry, e.g. photographic exposure meter using electric radiation detectors
- G01J1/44—Electric circuits
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
- G01J3/00—Spectrometry; Spectrophotometry; Monochromators; Measuring colours
- G01J3/02—Details
- G01J3/027—Control of working procedures of a spectrometer; Failure detection; Bandwidth calculation
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
- G01J3/00—Spectrometry; Spectrophotometry; Monochromators; Measuring colours
- G01J3/02—Details
- G01J3/10—Arrangements of light sources specially adapted for spectrometry or colorimetry
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
- G01J3/00—Spectrometry; Spectrophotometry; Monochromators; Measuring colours
- G01J3/28—Investigating the spectrum
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
- G01J3/00—Spectrometry; Spectrophotometry; Monochromators; Measuring colours
- G01J3/28—Investigating the spectrum
- G01J3/44—Raman spectrometry; Scattering spectrometry ; Fluorescence spectrometry
- G01J3/4406—Fluorescence spectrometry
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
- G01N21/17—Systems in which incident light is modified in accordance with the properties of the material investigated
- G01N21/47—Scattering, i.e. diffuse reflection
- G01N21/4738—Diffuse reflection, e.g. also for testing fluids, fibrous materials
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
- G01N21/62—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light
- G01N21/63—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light optically excited
- G01N21/64—Fluorescence; Phosphorescence
- G01N21/6408—Fluorescence; Phosphorescence with measurement of decay time, time resolved fluorescence
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
- G01N21/62—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light
- G01N21/63—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light optically excited
- G01N21/64—Fluorescence; Phosphorescence
- G01N21/645—Specially adapted constructive features of fluorimeters
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
- G01N21/62—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light
- G01N21/63—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light optically excited
- G01N21/64—Fluorescence; Phosphorescence
- G01N21/6486—Measuring fluorescence of biological material, e.g. DNA, RNA, cells
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N2201/00—Features of devices classified in G01N21/00
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- G01N2201/0625—Modulated LED
Abstract
method of monitoring a time-varying fluorescence emitted from a fluorescent agent from within a diffuse reflecting medium with time-varying optical properties is disclosed that includes providing at least two measurements obtained from a patient before and after administration of the fluorescent agent that includes anFlrsignal detected adjacent to the medium by a filtered light detector during illumination of the medium by excitatory-wavelength light, and at least one DR signal selected from: a and signal. The method further includes identifying a post-equilibration portion of the measurement data set and transforming each DRsignal within the post-equilibration portion of the measurement data set to an signal representing a detected fluorescence intensity emitted solely by the fluorescent agent from within the medium. The disclosed method includes removing the effects of leak-through of excitation-level light and removing the effects of autofluorescence from the DRsignal. gent that includes anFlrsignal detected adjacent to the medium by a filtered light detector during illumination of the medium by excitatory-wavelength light, and at least one DR signal selected from: a and signal. The method further includes identifying a post-equilibration portion of the measurement data set and transforming each DRsignal within the post-equilibration portion of the measurement data set to an signal representing a detected fluorescence intensity emitted solely by the fluorescent agent from within the medium. The disclosed method includes removing the effects of leak-through of excitation-level light and removing the effects of autofluorescence from the DRsignal.
Description
WO 40984 '1'
METHOD FOR VASIVE MONITORING OF FLUORESCENT TRACER
AGENT WITH BACKGROUND SEPARATION CORRECTIONS
CROSS-REFERENCE TO RELATED APPLICATION
This application claims the benefit of US. Provisional Application No.
62/452,021 filed January 30, 2017, which is incorporated herein in its ty.
BACKGROUND OF THE DISCLOSURE
The t disclosure relates generally to methods for non-invasive
monitoring of a fluorescent tracer agent within a medium characterized by scattering and/or
absorption of light. More particularly, the present disclosure relates to methods for non-
invasive assessment of kidney function by monitoring the clearance of an exogenous
fluorescent tracer within the tissues of a patient in vivo.
Dynamic monitoring of renal function in patients at the bedside in real time
is highly ble in order to minimize the risk of acute renal failure brought on by s
clinical, physiological and pathological conditions. It is particularly ant in the case of
critically ill or injured patients because a large percentage of these patients face the risk of
multiple organ failure (MOF) incited by one or more severe dysfunctions, such as: acute
lung injury (ALI), adult respiratory distress syndrome (ARDS), hypermetabolism,
nsion, persistent inflammation, and/or . Renal function may also be ed
due to kidney damage associated with administration of nephrotoxic drugs as part of a
ure such as angiography, diabetes, auto-immune disease, and other dysfunctions
and/or insults causally linked to kidney damage. In order to assess a t’s status and to
monitor the severity and/or progression of renal function over extended periods, there
exists considerable interest in ping a simple, accurate, and continuous method for the
determination of renal failure, preferably by non-invasive procedures.
Serum creatinine concentration, an endogenous marker of renal function, is
lly measured from a blood sample and used, in combination with patient
demographic factors such as weight, age, and/or ethnicity to estimate glomerular filtration
rate (GFR), one measure of renal function. However, creatinine-based assessments of renal
function may be prone to inaccuracies due to many potential factors, including: age, state
of hydration, renal perfusion, muscle mass, dietary intake, and many other anthropometric
and clinical variables. To compensate for these variances, a series of creatinine-based
equations (most recently extended to cystatin C) have been developed which incorporate
factors such as sex, race and other relevant factors for the estimation of glomerular
filtration rate (eGFR) based on serum creatinine ements. However, these eGFR
equations are not provided with any means of compensating for most of the above sources
of ce, and therefore have relatively poor accuracy. Further, the eGFR method
typically yields results that lag behind true GFR by up to 72 hrs.
Exogenous marker compounds, such as inulin, iothalamate, 51Cr-EDTA,
Gd-DTPA and 99mTc-DTPA have been used in existing methods for measuring GFR. Other
endogenous markers, such as 123I and 125I labeled o-iodohippurate or 99mTc-MAG3 have
been used to in other existing methods for assessing the tubular secretion process.
However, the use of typical exogenous marker nds may be accompanied by various
undesirable effects including the uction of ctive materials and/or ionizing
radiation into the patient, and laborious ex vivo handling of blood and urine samples,
rendering existing methods using these exogenous markers unsuitable for real-time
monitoring of renal on at a patient’s bedside.
The bility of a real-time, accurate, repeatable measure of renal
excretion rate using exogenous markers under patient-specific yet potentially changing
circumstances would represent a substantial improvement over any tly ced
method. Moreover, a method that depends solely on the renal elimination of an exogenous
al entity would provide a direct and continuous pharmacokinetic measurement
requiring less tive interpretation based upon age, muscle mass, blood pressure, etc.
BRIEF DESCRIPTION OF THE DRAWINGS
The disclosure will be better understood, and features, aspects and
advantages other than those set forth above will become nt when consideration is
given to the following detailed description thereof. Such detailed description makes
reference to the ing drawings, wherein:
is a schematic illustration of a single-wavelength renal monitoring
device in one aspect,
is a schematic illustration of a dual-wavelength renal monitoring
system in one aspect;
is a graph summarizing the absorption, transmission, and emission
spectra of various devices, materials, and compounds associated with the vasive
monitoring of an exogenous fluorescent agent in vivo defined over light wavelengths
ranging from about 430 nm to about 650 nm,
is a graph summarizing the absorption spectra of oglobin
(HbOz) and deoxyhemoglobin (Hb) defined over light wavelengths ranging from about 200
nm to about 650 nm,
is a schematic illustration of the timing of light pulse cycles
associated with data acquisition by a dual-wavelength renal monitoring system in one
aspect, in which each light pulse cycle includes light pulses produced at the tion
wavelength and at the on wavelength in sequence,
is a side view of a sensor head of a renal function ring system
in one aspect,
is a bottom view of the sensor head of
is a top interior view of the sensor head of illustrating an
ement of various electrical components within a housing of a sensor head of a renal
function monitoring system in one aspect,
is an enlargement of the interior view of
is a schematic ration of the apertures formed within a contact
surface of a sensor head of a renal function monitoring system in one aspect,
is a schematic illustration of synchronous detection of light by a
light detector of a sensor head in one aspect,
is a schematic illustration of light signal modulation and
demodulation by the sensor head in one aspect,
A is a block diagram illustrating the ts of a processing unit in
one aspect;
B is a block diagram illustrating the subunits of a processing unit in
a second aspect
is a graph of raw cence signal as a function of time
illustrating various phenomenon contributing to the total signal;
is a graph of intrinsic fluorescence signals, with and without an
autofluorescence correction, as a function of time illustrating the effect of an
autofluorescence correction on the renal decay time constants (RDTC) d from
analysis of the intrinsic fluorescence signal;
is a graph of raw fluorescence signal as a function of time in which
the final fluorescence signal falls below the original background fluorescence signal level
due to s phenomena contributing to the total signal;
A is a graph of raw fluorescence signal and excitation light leak-
through as a function of time;
B is a graph of raw fluorescence signal of A and a corrected
fluorescence signal comprising the raw fluorescence signal with the excitation light leak-
h of A removed;
is a graph comparing raw fluorescence signal (blue line) and
orescence signal (green line) obtained prior to injection of an exogenous fluorescent
agent;
A is a graph comparing raw fluorescence signal; autofluorescence
signal; and diffuse reflectance signals DRexmeas= DRem; and DRemflltgrgd ed prior to
ion of an exogenous fluorescent agent;
B is a graph comparing raw fluorescence signal; autofluorescence
signal; and diffuse reflectance signals DRexmeas= DRem; and DRemflltgrgd obtained after
injection of an exogenous fluorescent agent;
is a flow chart summarizing the steps of a background correction
method for removing the effects of excitation-wavelength light leak-through and
autofluorescence from the raw measured fluorescence signal;
is a graph of entative raw fluorescence signal measurements
(IFagem) detected by a renal monitoring device obtained before and after injection of an
exogenous fluorescent agent;
A is a block diagram illustrating a plurality of modules of a pre-
processing subunit in one aspect;
B is a block m illustrating a plurality of modules of a pre-
processing subunit in a second aspect;
is an isometric view of a sensor head of a renal on ring
system in a second ;
[003 5] is a bottom view of the sensor head of a renal function monitoring
system rated in ;
is an isometric view of the sensor head of a renal function
monitoring system illustrated in with the upper housing and various electrical
components removed to expose an inner housing;
is an exploded view of the inner housing of the sensor head
illustrated in ;
is a graph showing DRexmeasand Flrmeas over a full day in the
absence of administration of an ous fluorescent agent;
is a graph showing DRexmeasand Flrmeas immediately preceding
and following administration of an exogenous fluorescent agent; and
is a graph summarizing a relationship between empirically
DRem
determined Flrleakthmughand derived from a database of 33 ts.
DRemFilt
WO 40984 '6'
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 ming any incorporated
s. The patentable scope of the invention is defined by the claims, and may include
other examples that occur to those skilled in the art. Such other examples are intended to
be within the scope of the claims if they have structural elements that do not differ from the
l language of the claims, or if they include lent structural elements with
tantial differences from the literal languages of the claims.
DETAILED DESCRIPTION
Unless defined otherwise, all technical and scientific terms used herein have
the same meaning as commonly understood by one of ry skill in the art to which the
disclosure belongs. Although any methods and materials similar to or equivalent to those
described herein may be used in the practice or testing of the present disclosure, the
preferred materials and methods are described below.
A sample, as used herein, refers to a single, discrete data value acquired
from a signal and/or telemetry -to-digital converter (ADC) for a single
acquisition/telemetry channel.
A measured value, as used herein, refers to a single, discrete data value
created by demodulating or accumulating a sequence of samples from one acquisition
channel.
A measurement, as used herein, refers to a set comprising the Demodulated
In-Phase, lated Out-of—Phase, and Averaged measurement values from one
acquisition channel.
A measurement subset, as used herein, refers to a set comprising all
measurements for all ition channels during a single source LED nation. For
example, all measurements of an acquisition channel may include demodulated in-phase,
demodulated out-of-phase, and averaged measurements.
A measurement set, as used herein, refers to a set comprising one
measurement subset for each source LED.
'7' 2018/016053
An acquisition, as used herein, refers to the overall s by which a
measurement set is obtained.
A measurement sequence, as used herein, refers to a sequence of one or
more measurement sets.
A telemetry value, as used herein, refers to a single, discrete data value
acquired from a single channel of a telemetry ADC.
A try set, as used herein, refers to a set comprising one telemetry
value from each telemetry channel.
A diffuse reflecting medium, as used herein, refers to any material through
which light propagates, which includes a plurality of moieties, particles, or molecules that
may scatter, reflect, and/or absorb the light as it propagates. The bution of the
plurality of moieties, particles, and/or molecules may be m or non-uniform, and may
change over time.
In various s, systems and methods for monitoring time-varying
fluorescence d from a fluorescent agent from within a diffuse reflecting medium with
time-varying optical properties are disclosed herein below. In one aspect, systems and
methods for monitoring a time-varying fluorescence emitted from an exogenous
fluorescent agent within the tissues of a patient are disclosed. The systems and methods of
this one aspect may be used in a variety of contexts including, but not limited to, the
monitoring of renal function in vivo in a patient in real time by monitoring the decreasing
fluorescence emitted by an exogenous fluorescent agent within the tissue of a patient as the
exogenous fluorescent agent is eliminated by the kidneys of the patient. Although the
systems and devices disclosed herein below are described in the context of methods and
devices to monitor kidney function, it is to be understood that the sed systems and
methods may be applied to any s and methods that r the arying
fluorescence emitted by a fluorescent agent from within a diffuse reflecting , in
which the optical ties of the diffuse reflecting medium may also vary with time.
is a schematic illustration of a system 100, provided as a non-limiting
example, in which fluorescence 102 with an emission wavelength (hem) is detected from a
region of interest of a t 104 using a light detector 110 configured to detect only those
photons with an emission wavelength (hem). In general, the exogenous fluorescent agent
112 produces fluorescence 102 in se to an excitation event ing, but not limited
to: nation by light 106 at an excitation wavelength (hex), occurrence of an enzymatic
reaction, changes in local ical potential, and any other known excitation event
associated with exogenous fluorescent agents. In an aspect, the system 100 may include a
light source 108 configured to deliver light 106 at an excitation wavelength (hex) to the
patient 104. In this aspect, the fluorescence 102 is produced in response to illumination by
the light 106. In addition, the excitation wavelength (hex) of the light 106 and the emission
wavelength (hem) of the fluorescence 102 are ally ct (i.e., hex is sufficiently
ent from hem) so that the light detector 110 may be configured to selectively detect
only the fluorescence 102 by the inclusion of any known optical wavelength separation
device including, but not d to, an optical filter.
In some s, changes in the fluorescence 102 may be monitored to
obtain information regarding a physiological function or status of the patient. By way of
non-limiting example, the time-dependent decrease in the fluorescence 102 ed after
introduction of the exogenous fluorescent agent 112 into a circulatory vessel of the patient
104 may be analyzed to obtain information regarding renal function of the t 104. In
this non-limiting example, the rate of decrease in cence 102 may be assumed to be
proportional to the rate of l of the exogenous fluorescent agent 112 by the kidneys
of the patient 104, thereby providing a measurement of renal function including, but not
limited to: renal decay time constant (RDTC) and glomerular filtration rate (GFR).
Without being limited to any particular theory, the intensity of fluorescence
102 detected by the light detector 110 may be influenced by any one or more of numerous
factors including, but not limited to: the intensity or power of the light 106 at hex delivered
to the patient 104, the scattering and absorption of the light 106 passing through
intervening tissues 114 of the patient 104 between the light source 108 and the exogenous
fluorescent agents 112, the concentration of exogenous fluorescent agents 112 illuminated
by the light 106, the scattering and absorption of the fluorescence 102 at hem passing
through intervening tissues 114 of the patient 104 n the exogenous fluorescent
agents 112 and the light detector 110, leak-through of the excitation light 106 through any
optical filters configured to transmit only light at emission wavelength hem, and
fluorescence emitted by endogenous skin ents.
is a graph showing a representative time y of a raw
fluorescence signal obtained at the emission wavelength hem corresponding to the
wavelength of cence emitted by an endogenous fluorescent agent within the tissues
of a patient in response to illumination by tion-wavelength light. The measured raw
fluorescence signal obtained prior to the injection of the endogenous fluorescence agent
(i.e. the background signal 1402) may include autofluorescence (Fauw) emitted by
endogenous ures as well as leak-through of tion-wavelength light (ExLI)
through any optical filters configured to transmit only emission-wavelength light to the
light detector producing the raw cence signal. The measured raw fluorescence
signal obtained after the ion of the endogenous fluorescent agent (i.e. Flrmeas 1404)
may include the ity of the fluorescence emitted by the endogenous fluorescent agent
(Fagem) superimposed over the background signal 1402 (i.e. FWD and ExLI).
Existing methods typically assume that the optical properties within the
intervening tissue 114 remain essentially unchanged throughout the period during which
ements are obtained by the system 100. As a result, existing methods typically
obtain l measurements through the intervening tissue 114 of the patient 104 prior to
introduction of the exogenous fluorescent agent 112, and these initial measurements are
subtracted to correct all subsequent data obtained after 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 ing, but not limited to: optical coupling efficiency of the light
detector 110 to the patient 104, concentration of chromophores such as hemoglobin due to
changes in blood volume caused by vascular dilation, iction, or compression,
changes in the optical properties of chromophores such as hemoglobin due to changes in
oxygenation status, and changes in tissue structure such as changes related to edema.
is a graph of the raw fluorescence signal measured before and after
the injection of an endogenous fluorescent agent, illustrating that the background signal
may change over the extending data ition period associated with the renal clearance
of the endogenous fluorescent agent from the patient. As illustrated in , the initial
W0 2018/140984 PCT/U82018/016053
background signal level 1602 is about 0.01 intensity units higher than the final background
signal level 1604 measured about nine hours after the measurement of the initial
background signal level 1602. Without being limited to any particular theory, it is thought
the administration of a blood pressure medication during the data ition period may
have induced skin flushing and associated vasodilation of skin aries that may have
altered the l properties of the patient’s skin, due to the increased concentration of
blood, which contains hemoglobin, an known endogenous chromophore capable of
absorbing light at both the excitation and emission wavelengths.
These dynamic changes in the optical properties of the ening tissue
114 may introduce uncertainty into long-term measurements of fluorescence 102. By way
of miting example, changes in the optical properties of the intervening tissue 114
may modulate the intensity or power of the light 106 illuminating the exogenous
fluorescent agents 112, causing a modulation of the fluorescence 102 produced by the
exogenous fluorescent agents 112 that may be erroneously interpreted as a tion in
the tration of the ous cent agents 112. By way of r miting
example, changes in the optical properties of the intervening tissue 114 may modulate the
intensity or power of the fluorescence 102 reaching the light detector 110 that may also be
erroneously reted as a modulation in the concentration of the exogenous fluorescent
agents 112. The potential modulation of changes in the optical properties of the intervening
tissue 114 may uce ainty into measurements of fluorescence 102, in particular
those measurements associated with long-term monitoring of fluorescence 102 as described
herein above.
Similarly, because autofluorescence (Fame) produced by endogenous
chromophores occurs in a similar manner to the fluorescence produced by the exogenous
fluorescent agent, dynamic changes in the optical properties of the intervening tissue may
introduce variability in the autofluorescence (Fame) levels over the course of long-term
measurements of fluorescence 102. By way of non-limiting example, changes in the
scattering and absorption of the light 106 passing through the intervening tissue 114 may
modulate the intensity or power of the light 106 illuminating the endogenous
chromophores, causing a modulation of the autofluorescence that may modulate the
background fluorescence over the course of data acquisition. By way of another non-
limiting example, changes in the scattering and absorption of the autofluorescence passing
W0 2018/140984 PCT/U82018/016053
through the intervening tissue 114 may modulate the intensity of the autofluorescence
detected by the light detector 110 that may te the background fluorescence over the
course of data acquisition. The potential modulation of background cence, if not
ly accounted for, may introduce uncertainty into raw fluorescence measurements and
by extension may introduce uncertainty into parameters derived from an analysis of these
florescence measurements.
By way of non-limiting example, changes in autofluorescence d to
dynamic changes in the optical properties of the skin of the patient may introduce
uncertainty into the calculation of renal decay time constant (RDTC), a measure of renal
function as bed herein below. is a graph of a raw fluorescence signal
measured before and after the injection of an endogenous fluorescent agent that includes
autofluorescence (IFAgent+Aut0F;r, blue line). The graph of also includes a corrected
fluorescence signal m, green line) calculated by removing the effects of
autofluorescence from the raw fluorescence signal using the methods described herein
below. Superimposed on each signal are flts associated with the calculation of
RDTC. As shown in , the RDTC value of 2.76 hr. ated using the raw
fluorescence signal is considerably higher that the corresponding RDTC value of 2.31 hr.
calculated using the corrected fluorescence signal.
In various aspects, a method of correcting in vivo real-time measurements of
fluorescence from an exogenous fluorescent agent to remove the effects of changes in the
optical properties within the tissue of the patient is provided. The inclusion of an
additional measurement of light passing through the tissue of the patient via a separate
optical pathway (i.e. diffuse reflectance) from the optical pathway of the fluorescence
measurements enhanced the quantification of changes in the optical properties of the tissue
during prolonged monitoring of fluorescence from an exogenous fluorescent agent within a
patient. The inclusion of this additional measurement in the correction method in s
s was discovered to significantly enhance the y of cence measurements.
Detailed descriptions of devices for monitoring the fluorescence of an
ous fluorescent agent in vivo and methods of correcting the fluorescence
measurements to remove the effects of changes in the background signal are provided
herein below.
W0 2018/140984 PCT/U82018/016053
Although the devices and methods are described herein below in the t
of a non-invasive optical renal function monitor, it is to be understood that the correction
method described herein, with appropriate modification, may be applied to any compatible
device configured to m ements by delivering EM radiation from an external
source through any scattering medium and/or ing EM radiation propagated h
any scattering medium to an external detector. Non-limiting examples of EM radiation
include visible light, near-IR light, IR light, UV radiation, and microwave radiation. The
scattering media may include any living or non-living material capable of propagating EM
radiation of at least one EM frequency without tion. At least a portion of the
scattering media may further include one or more substructures or compounds capable of
ing and/or absorbing the EM radiation. Non-limiting examples of scattering media
include: a tissue of a living or dead organism, such as a skin of a mammal, a gas such as air
with or without additional particles such as dust, fluid droplets, or a solid particulate
material, a fluid such as water with or without onal particles such as gas bubbles or a
solid particulate material. r, the devices and methods described herein below are not
limited to detection of renal function, but may be modified for use in the detection of the
function of other physiological systems including, but not limited to, liver systems, or
gastro-intestinal s.
System Description
In various aspects, the methods of correcting fluorescence measurements to
remove the s of variations in local skin properties as disclosed herein may be
orated into any fluorescence monitoring system including, but not limited to, a
system for optically monitoring renal function in vivo and in real time by ing
changes in fluorescence of an exogenous fluorescent agent injected into a patient as the
agent is renally eliminated from the patient. is a block diagram of a system 200 for
optically monitoring renal function of a patient 202 via measurements of the fluorescence
of an injected exogenous fluorescent agent in the patient 202, in one aspect. The system
200 may include at least one sensor head 204 configured to deliver light at an excitatory
wavelength (hex) into a first region 206 of the patient 202. The system 200 is further
configured to detect light at an emission wavelength (hem), at a second region 208 of the
patient 202, and to detect light at the excitatory wavelength (hex), and/or emission
wavelength (hem), at a third region 210 of the patient 202.
W0 2018/140984 PCT/U82018/016053
The system 200 may further include a controller 212 operatively d to
the at least one sensor head 204, an ion unit 214, and a display unit 216. In various
s, the controller 212 is red to control the operation of the at least one sensor
head 204 as described in additional detail herein below. The controller 212 is further
configured to receive measurements of light from the at least one sensor head 204. The
controller 212 is further configured to correct the light measurements corresponding to
fluorescence from exogenous fluorescent agents according to at least one method
including, but not limited to, the disclosed methods of correcting fluorescence
measurements using measurements indicative of dynamic changes in the background signal
related to changes in autofluorescence and/or the leak-through of excitatory-wavelength
light to the second light detector 224 configured to detect on-wavelength light only.
The controller 212 is further configured to transform the fluorescence measurements
received from the at least one sensor head 204 into a summary parameter representative of
the renal function of the t 202. In addition, the controller 212 is red to receive
at least one signal representing user inputs from the operation unit 214 and to generate one
or more forms for display on the display unit 216 including, but not limited to, a graphical
user ace (GUI).
A detailed description of the sensor head 204 and controller 212 are
provided herein below.
A. Sensor Head
In various aspects, the sensor head 204 includes at least one light source and
at least one light detector in a housing. is a side view of a g 600 for the
sensor head 204 in one aspect that includes an upper housing 602 and a lower housing 604
attached together to enclose two light sources and two light detectors. The bottom surface
608 of the lower housing 604 further includes a contact surface 606 configured to be
ed to the skin of a patient 202 using a patible adhesive material including, but
not limited to, a surgical adhesive. In use, the surface of the adhesive material opposite to
the contact surface 606 may be affixed to the skin of the patient 202. In various aspects,
the adhesive material may be configured to transmit light through the light s into the
patient and to further transmit the fluorescence from the patient to the light detectors. In
one aspect, the adhesive material may be an optically transparent al. In another
W0 2018/140984
aspect, the ve material may be produced from a non-fluorescing material to prevent
the production of confounding fluorescence by the adhesive material.
In various other aspects, the upper g 602 may r include one or
more openings 806 configured to provide access to the interior for a cable including, but
not limited to, a USB cable, and/or to e a window for a y generated by the
circuitry contained within the housing 600, such as an indicator LED.
is a bottom view of the housing 600 illustrated in The
contact surface 606 may include an aperture plate 702 including one or more apertures 704
configured to transmit light between the skin of the t and the light s and light
detectors contained inside the housing 600. In one aspect, the aperture plate 702 may be
epoxied into the lower housing 604 to prevent liquid ingress into the interior of the housing
600. In various aspects, the dimensions, arrangement, and/or spacing of the one or more
apertures 704 may be selected to enhance various aspects of the operation of the system
200, as described in additional detail herein below. In another , the t surface
606 may further include a temperature sensor opening 706 configured to provide a thermal
path from the skin surface of the t to an additional temperature sensor 228 configured
to monitor the temperature at the skin surface of the patient.
is a schematic m illustrating the arrangement of the electrical
components within the housing 600. Referring to the upper housing 602 and the
lower housing 604 may be affixed together with screws 802, and the screw holes and the
interface between the two housing pieces may be filled with a resistant filler material
804 including, but not limited to, a silicone material such as room temperature
vulcanization silicone (RTV) to inhibit liquid ingress into the interior of the housing 600.
In an aspect, the housing 600 may further include a cable opening 806
formed through the upper housing 602. The cable opening 806 may be configured to
provide access to the interior for an electrical cable including, but not limited to, a USB
cable. In one , the cable may enable the supply of power to the light sources, light
detectors, tor lights, and associated electrical devices and circuits as described herein
below. In another , the cable may further enable the communication of control
signals into the housing to enable the operation of the electrical components within the
housing 600, and the cable may further enable the communication of data signals encoding
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measurements obtained by one of more of the sensor devices contained within the housing
600 ing, but not limited to: the first light detector 222, the second light detector 224,
any additional light detectors, such as a first monitor iode 904 and a second monitor
diode 906, and any additional temperature sensors 228 (see . In an aspect, the cable
may be attached to the cable opening 806 and adjacent upper housing 602 with a light
absorbent adhesive including, but not limited to, black epoxy and may further be sealed
against water incursion using a water resistant filler material including, but not limited to,
RTV.
In an additional aspect, the housing 600 may further include at least one
display opening 808 formed through the upper housing 602. In one aspect, each display
opening 808 may be configured to e a window for a display generated by the
try contained within the housing 600, such as an indicator LED 810. In an aspect,
each indicator LED 810 may be oned on a circuit board 812. In an aspect, a light pipe
814 may be epoxied into the display opening 808 within the upper housing 602 above each
indicator LED 810. Each a light pipe 814 may be filled with a water-resistant filler material
such as RTV for liquid ingress protection. In s aspects, the at least one tor LED
810 may illuminate in a predetermined pattern to enable a user of the system 200 to
r the operational status of the sensor head 204.
is a close-up view of the interior optical region of the sensor head
204 g the arrangement of the light sources 2l8/220 and the light detectors 222/224
within the housing 600 in one aspect. In an aspect, the light sources 2l8/220 are separated
from the light detectors 222/224, and the first light detector 222 is separated from the
second light detector 224 are separated from one another by a sensor mount 9l2 affixed to
the aperture plate 702. In an aspect, the sensor mount 9l2 ensures that light from the light
sources 2l8/220 does not reach the light detectors 222/224 without coupling h the
skin of the t 202. The separation between the first light detector 222 within the first
detection well 908 and the second light detector 224 within the second detection well 910
ensures that the fluorescence signal produced by the exogenous fiuorescent agent within
the tissues of the patient 202 is distinguishable from the unfiltered excitation light
introduced by the first light source 218.
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Referring again to the sensor mount 912 may be aligned to a circuit
board (not shown) containing the light sources 0 and light detectors 222/224 using
alignment pins 914 and held in place using screws 916. In an aspect, the sensor mount 912
may be affixed to the circuit board containing the light s 218/220 and light detectors
222/224 using a light absorbent adhesive including, but not d to, black epoxy. In this
aspect, this light-resistant join between the circuit board and the sensor mount 912 inhibits
leakage of light between the light sources 218/220 and the light detectors 222/224, and
further inhibits the leakage of light between the first light detector 222 and the second light
detector 224. The apertures 704 configured to transmit light to and from the skin
underlying the contact surface 606 of the sensor head 204 are formed through a structurally
te aperture plate 702 (see to provide for precise alignment of the apertures
704 to the corresponding light s 0 and light detectors 222/224, described in
additional detail herein below.
In various aspects, the sensor mount 912 may further provide electrical
shielding for any sensitive electrical devices within the sensor head 204 ing, but not
limited to, the light detectors 222/224. In one aspect, the sensor mount 912 may be
constructed of an electrically conductive material including, but not limited to: aluminum
and aluminum alloy. In this aspect, the sensor mount 912 may be electrically coupled to the
ground of the circuit board using conductive screws 916. In on, any glass windows
positioned within the source well 902 and/or or wells 908/910 adjacent to the
aperture plate 702 including, but not limited to, an optical filter 244 and clear glass 246 as
described herein below (see may further include an electrically conductive coating.
Non-limiting examples of suitable electrically conductive coatings for the glass windows of
the sensor mount include a conductive indium tin oxide (ITO) coating and any other
suitable transparent and electrically conductive coating.
Without being d to any particular theory, the conductive material of
the sensor mount 912 provides a l Faraday cage to shield the electrically sensitive
detectors 222/224 from electrical noise generated by or conducted through the patient’s
body. The partial Faraday cage ed by the sensor mount 912 may be completed with
the conductive ITO g on the glass windows within the source well 902 and/or
detector wells 908/910. In an aspect, the electrically conductive coating on the glass
windows, such as an ITO coating, are sufficiently conductive to provide ical shielding
W0 2018/140984
while remaining sufficiently transparent for the transmission of light to and from the skin
surface of the patient 202. In another , the ITO coating of each glass window may be
grounded to an ically conductive sensor mount 912 using any known electrical
grounding method including, by not limited to: a wire ting the glass coating to the
sensor mount 912 that is attached at both wire ends with conductive epoxy, or attaching the
coated glass directly to a glass fitting such as a ledge or frame formed within each of the
source well 902 and/or detector wells 908/910 using an electrically conductive epoxy.
In various aspects, the contact surface 606 of the g 600 may be
attached the patient’s skin using a patible and an adhesive material 610 ing,
but not limited to, a clear double-sided medical grade adhesive, as illustrated in and
Any adhesive material selected to be optically transmissive at the excitation and
emission wavelengths used by the system 100 as described herein. The adhesive material
610 may be positioned on the contact surface 606 such that the adhesive material covers
the apertures 704, but exposes the temperature sensor opening 706 to ensure sufficient
thermal contact with the skin of the patient 202. In one aspect, the sensor head 204 may be
further secured to the patient 202 as needed using one or more additional patible
medical fastener devices including, but not limited to: Tegaderm bandages, medical tape,
or any other suitable biocompatible medical er devices.
In an aspect, the contact surface 606 may be d near the front edge of
the sensor head 204 to provide for te positioning of the contact surface 606 on a
ed region of the patient’s skin. In another aspect, the apertures 704 may be positioned
towards the center of the contact surface 606 to reduce ambient light ingress. Without
being limited to any particular , ambient light may enter one or more of the apertures
704 due to incomplete on of the contact surface 606 to the patient’s skin and/or due
to the propagation of ambient light passing through the patient’s exposed skin situated just
outside of the footprint of the contact surface 606 into the apertures 704.
Referring again to the bottom surface 608 of the sensor head 204
curves away from the plane of the contact surface 606 to enable attachment of the sensor
head 204 to varied body type and locations. For attachment of the sensor head 204 to
relatively flat or concave surfaces, any gap 612 between the bottom surface 608 and the
W0 2018/140984
skin surface of the patient 202 may be filled with a biocompatible foam to ensure
consistent contact with the patient 202.
1) Light sources
In various aspects, each sensor head 204 includes a first light source 218
and a second light source 220 configured to r light to a first region 206 of a patient
202. The first light source 218 is configured to deliver the light at the excitatory
wavelength and the second light source 220 is configured to r light at the emission
ngth. In one aspect, the excitatory wavelength may be selected to fall within a
spectral range at which the exogenous cent agent ts relatively high
absorbance. In another aspect, the on wavelength may be selected to fall within a
spectral range at which the exogenous fluorescent agent exhibits relatively high emission.
The exogenous fluorescent agent may be selected for enhanced contrast ve to other
chromophores within the tissues 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 may be selected to conduct measurements within spectral
ranges with lower variation in absorption by other chromophores such as hemoglobin
within the tissues of the patient 202 during use.
Without being limited to any particular theory, hemoglobin (Hb) is an
absorber of visible light in the tissues of the patient 202, and has the potential to interfere
with the measurements of fluorescence of the exogenous fluorescent agent if the Hb
absorbance varies over the measurement period of the system 200. Because hemoglobin
(Hb) enables gas exchange within virtually all tissues containing circulatory vessels,
virtually all tissues are vulnerable to interference with fluorescence measurements of the
system 200 due to fluctuations in hemoglobin tration. Within most tissues,
externally applied pressure may cause blood g which may be manifested as an
apparent decay of the fluorescence measured at the skin surface. Periodic opening and
closing of blood vessels motion”) near the surface of the skin may also cause
fluctuations in hemoglobin concentration which may introduce additional noise in to
ements of fluorescence of the exogenous cent agent by the system 200.
r, in some patients 202, such as those with pulmonary disorders, variation in the Hb
oxygenation state may also be observed, leading to additional potential variations in the
W0 2018/140984
background skin absorbance due to differences in the absorption spectra of
deoxyhemoglobin (Hb) and oxyhemoglobin (HbOz), shown rated in
In an aspect, the excitation and emission wavelengths for the exogenous
fluorescent agent may be ed to coincide with a pair of HbOz/Hb stic points,
each isosbestic point defined herein as a wavelength characterized by about equal light
absorbance by HbOz and Hb. Without being limited to any ular theory, fluorescence
measurements conducted at each isosbestic wavelength are less sensitive to variation due to
changes in the oxygenation of obin, so long as the ed concentration of HbOz
and Hb remains relatively stable during measurements of fluorescence by the system 200.
Non-limiting examples of Hb/HbOz 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.
In various aspects, the excitation and emission wavelengths may be ed
based on the tion and emission wavelengths of the selected exogenous fluorescent
agent of the system 200. In one aspect, the excitatory wavelength may be an HbOz/Hb
isosbestic wavelength and simultaneously may be a wavelength within a spectral range of
high absorbance of the exogenous fluorescent agent. In another aspect, the emission
wavelength may be an HbOz/Hb isosbestic wavelength and simultaneously may be a
ngth within a spectral range of emission by the exogenous fluorescent agent. Table 1
provides a summary of b isosbestic wavelengths within the spectral range of 200
nm to about 1000 nm. is a graph of the absorption spectra used to identify the
HbOz/Hb isosbestic wavelengths of Table 1.
Table I. Hb02/Hb Isosbesll'c Wavelengths )t = 200 - 1000 nm
Excitation Hb Molar Hb02 dA/d)» Hb dA/dl
Wavelength Extinct. Coeff. (M'1 cm'1 nm'l)
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By way of rative example, is a graph izing the
absorption a for HbOz and Hb, as well as the absorption and on spectra of
frequency spectra of 1Vfl3402, an exogenous fluorescent agent in one aspect. Emission
spectra for a blue LED light source and a green LED light source are also shown
superimposed over the other spectra of In this , the system 200 may include
a blue LED as the first light source 218, and the excitatory ngth for the system 200
may be the isosbestic wavelength of about 450 nm. As listed in Table 1 and shown in the Hb absorbance spectra is strongly sloped at the isosbestic wavelengths of about 420
nm to about 450 nm (see columns 3 and 4 of Table 1), indicating that the relative
absorbance of HbOz and Hb at the isosbestic wavelength of about 450 nm is sensitive to
small changes in excitatory wavelength. However, at wavelengths above about 500 nm, the
HbOz/Hb spectra are less steeply sloped, and a broader band light source including, but not
d to, an LED with a bandpass filter may suffice for use as a first light source 218.
In another aspect, the excitatory wave length may be selected to enhance the
contrast in light absorbance between the exogenous fluorescent agent and the
chromophores within the tissues of the patient 202. By way of non-limiting example, as
shown in at the isosbestic ngth of 452 nm, the light absorption of the MB-
102 is more than three-fold higher than the light absorption of the HbOz and the Hb.
Without being limited to any particular theory, a higher proportion of light illuminating the
tissue of the patient 202 at a wavelength of about 450 nm will be absorbed by the 1Vfl3402
relative to the HbOz and Hb, thus enhancing the efficiency of absorption by the 1Vfl3402
and reducing the intensity of light at the excitatory wavelength needed to elicit a detectable
fluorescence signal.
In various aspects, a second isosbestic wavelength may also be selected as
the emission wavelength for the system 200. By way of miting example,
shows an emission spectrum of the MB402 exogenous contrast agent that is characterized
W0 2018/140984 PCT/U82018/016053
by an emission peak at a wavelength of about 550 nm. In this non-limiting example, the
isosbestic wavelength of 570 nm may be selected as the emission ngth to be
detected by first and second detectors 222/224. In various other aspects, the emission
wavelength of the system 200 may be ed to fall within a spectral range characterized
by vely low absorbance of the chromophores within the tissues of the patient 202.
Without being d to any particular theory, the low absorbance of the chromophores at
the selected emission wavelength may reduce the losses of light emitted by the exogenous
fluorescent agent and enhancing the efficiency of fluorescence detection.
In various aspects, the first light source 218 and the second light source 220
may be any light source configured to deliver light at the excitatory wavelength and at the
emission wavelength. Typically, the first light source 218 delivers light at an ity that
is sufficient to penetrate the s of the t 202 to the exogenous fluorescent agent
with sufficient intensity remaining to induce the emission of light at the emission wave
length by the exogenous fluorescent agent. Typically, the first light source 218 rs
light at an intensity that is sufficient to penetrate the tissues of the patient 202 to the
exogenous fluorescent agent with sufficient intensity remaining after scattering and/or
absorption to induce fluorescence at the emission wave length by the exogenous
fluorescent agent. However, the intensity of light delivered by the first light source 218 is
limited to an upper value to prevent adverse effects such as tissue burning, cell damage,
and/or photo-bleaching of the exogenous fluorescent agent and/or the endogenous
chromophores in the skin (“auto-fluorescence”).
rly, the second light source 220 rs light at the emission
wavelength of the exogenous fluorescent agent at an intensity red to provide
sufficient energy to propagate with scattering and tion through the first region 206 of
the patient and out the second region 208 and third region 210 with sufficient remaining
intensity for detection by the first light detector 222 and the second light detector 224,
respectively. As with the first light source 218, the intensity of light ed by the
second light source 220 is limited to an upper value to prevent the adverse effects such as
tissue injury or photobleaching described previously.
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In various aspects, the first light source 218 and the second light source 220
may be any light source suitable for use with fluorescent medical imaging systems and
devices. Non-limiting examples of suitable light sources include: LEDs, diode lasers,
pulsed lasers, continuous waver lasers, xenon arc lamps or mercury-vapor lamps with an
excitation filter, lasers, and supercontinuum sources. In one aspect, the first light source
218 and/or the second light source 220 may produce light at a narrow spectral bandwidth
suitable for monitoring the concentration of the exogenous fluorescence agent using the
method described herein. In another , the first light source 218 and the second light
source 220 may produce light at a relatively wide spectral bandwidth.
In one , the selection of intensity of the light produced by the first
light source 218 and the second light source 220 by the system 200 may be influenced any
one or more of at least several factors including, but not limited to, the maximum
permissible exposure (MPE) for skin exposure to a laser beam according to applicable
regulatory standards such as ANSI rd Z1361 In another aspect, light intensity for
the system 200 may be selected to reduce the likelihood of photobleaching of the
exogenous fluorescent source and/or other chromophores within the tissues 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 may be selected in order to elicit a detectable fluorescence signal from the
exogenous fluorescent source within the tissues of the t 202 and the first light
detector 222 and/or second light detector. In yet r aspect, the light intensity for the
system 200 may be selected to provide ly high light energy while ng power
consumption, inhibiting heating/overheating of the first light source 218 and the second
light source 220, and/or reducing the exposure time of the patient’s skin to light from the
first light detector 222 and/or second light detector.
In s aspects, the intensity of the first light source 218 and the second
light source 220 may be modulated to compensate any one or more of at least several
factors including, but not limited to: individual differences in the tration of
phores within the patient 202, such as variation in skin pigmentation. In various
other aspects, the detection gain of the light detectors may be modulated to similarly
sate for variation in individual differences in skin properties. In an aspect, the
variation in skin pigmentation may be between two different individual patients 202, or
W0 2018/140984 PCT/U82018/016053
between two different ons on the same patient 202. In an aspect, the light modulation
may compensate for variation in the optical pathway taken by the light through the tissues
of the patient 202. The optical pathway may vary due to any one or more of at least several
factors including but not limited to: variation in separation distances between the light
sources and light detectors of the system 200; variation in the secure attachment of the
sensor head 204 to the skin of the patient 202; variation in the light output of the light
sources due to the exposure of the light sources to environmental factors such as heat and
moisture; variation in the sensitivity of the light detectors due to the exposure of the light
ors to environmental factors such as heat and moisture; modulation of the duration of
illumination by the light sources; and any other relevant operational ter.
In various aspects; the first light source 218 and the second light source 220
may be configured to modulate the ity of the light produced as needed according to
any one or more of the factors described herein above. In one aspect; if the first light
source 218 and the second light source 220 are devices configured to continuously vary
output fluence as needed; for example LED light s; the intensity of the light may be
modulated onically using methods including; but not d to; modulation of the
electrical 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 may be ted
using optical s including; but not limited to: partially or fully occluding the light
leaving the first light source 218 and the second light source 220 using an optical device
including; but not limited to: an iris; a shutter; and/or one or more ; diverting the path
of the light leaving the first light source 218 and the second light source 220 away from the
first region 206 of the patient using an optical device including; but not limited to a ;
a mirror; and/or a prism.
In various s; the intensity of the light produced by the first light source
218 and the second light source 220 may be modulated via control of the laser fluence;
defined herein as the rate of energy within the produced light beam. In one aspect; the laser
fluence may be d to ranges defined by safety standards including; but not limited to;
ANSI standards for exposure to laser energy such as ANSI Zl36.l. Without being limited
to any particular theory; the maximum fluence of light delivered to a patient 202 may be
influenced by a variety of factors including; but not limited to the wavelength of the
delivered light and the duration of exposure to the light. In various aspects; the maximum
W0 2018/140984
fluence of light may range from about 0.003 J/cm2 for light at delivered at wavelengths of
less than about 302 nm to about 1 J/cm2 for light delivered at wavelengths ranging from
about 1500 nm to about 1800 nm for a duration of up to about 10 sec. For light delivered
at wavelengths ranging from about 400 nm to about 1400 nm (visible/NIR light) the
maximum fluence may be about 0.6 J/cm2 for a duration of up to about 10 sec, and up to
about 0.2 J/cm2 for a duration ranging from about 10 sec to about 30,000 sec. For
extended exposures, the delivered light is limited to a maximum power density (W/cm2)
ing to ANSI standards: visible/NIR light is limited to 0.2 W/cm2 and far IR light is
limited to about 0.1 W/cm2. Without being limited to a particular theory, ed
exposure to light delivered at UV wavelengths is not typically recommended according to
ANSI standards.
In another aspect, the fluence of light at the excitatory wavelength produced
by the first light source 218 may be modulated in order to provide sufficient energy to
propagate through the skin in the first region 206 of the patient 202 to the exogenous
fluorescent agent without photobleaching, and to nate the exogenous fluorescent
agent with energy sufficient to induce able fluorescence at the first light detector 222
and/or the second light detector 224. In an additional aspect, the fluence of light at the
emission wavelength produced by the second light source 220 may be modulated in order
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 t
photobleaching to emerge as detectable light at the first light detector 222 and the second
light or 224, respectively. By way of non-limiting example, the fluence of light
produced by a light source at 450 nm or 500 nm may be limited to 1.5 and 5 mW/cm2,
tively, to prevent photo-bleaching.
In s aspects, the fluence of the light produced by the first light source
218 and the second light source 220 may be modulated by any suitable systems and/or
s without limitation as described herein above. This modulation may be enabled a
single time during operation of the system 200, and as a result, the fluence of the light
produced by each of the first light source 218 and the second light source 220 may be
relatively constant throughout the operation of the system 200. In another aspect, the light
modulation may be enabled at te times over the duration of operation of the system
W0 2018/140984 PCT/U82018/016053
200, or the light tion may be enabled continuously over the duration of operation of
the system 200.
In one aspect, the fluence of the light may be modulated via manual
adjustment of any of the power source settings and/or optical device settings as described
above when the system 200 is configured in an Engineering Mode. In another aspect, the
fluence of the light may be modulated tically via one or more control schemes
encoded in the light source control unit of the controller 212 as described herein below. In
this , the degree of modulation may be ed at least in part on the basis of
feedback measurements obtained by various sensors provide in the sensor head 204 of the
system 200 including, but not limited to, additional light detectors 226 and temperature
sensors 228 as described in additional detail herein below.
In s aspects, light produced by the first light source 218 and the
second light source 220 are further characterized by a pulse width, defined herein as the
duration of the produced light. Although pulse width is typically used to characterize the
performance of a light source that produces light in te pulses, such as a pulsed laser,
it is to be understood that the term “light , as used herein, refers to any discrete burst
of light produced by a single light source at a single wavelength to enable the acquisition of
a single measurement of fluorescence 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 ing, but not limited to: delivery of sufficient light energy to elicit detectable
fluorescence from the exogenous fluorescent agent without photobleaching the exogenous
fluorescent agent or other chromophores within the tissues of the t 202, compliance
with safety standards for light ry to patients such as ANSI rds, light delivery at
sufficiently high rate to enable data acquisition 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 working life of light
sources, light detectors, and other devices related to producing and detecting light energy,
and any other relevant factors.
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In various aspects, the pulse width of the light produced by the first light
source 218 and the second light source 220 may be independently selected to be a on
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 may be ndently selected to be a duration ranging from about 0.0001 seconds to
about 0.001 s, from about 0.0005 seconds to about 0.005 s, 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 s to about
0.35 seconds, from about 0.3 seconds to about 0.4 seconds, from about 0.35 seconds to
about 0.45 seconds, and from about 0.4 seconds to about 0.5 seconds. In one aspect, the
pulse widths of the light produced by the first light source 218 and the second light source
220 are both about 0.1 seconds, as illustrated schematically in
In another , the light produced by the first light source 218 and the
second light source 220 may be further characterized by a pulse rate, defined herein as the
number of pulses produced by a light source per second. Although pulse rate is typically
used to characterize the performance of a light source that produces light in discrete pulses,
such as a pulsed laser, it is to be understood that the term “pulse rate”, as used herein,
refers to the rate of production of a discrete light pulse by a single light source at a single
wavelength in association with the ition of measurements of fluorescence by the
system 200. In various aspects, the pulse rate may be selected based on one or more of at
least several factors including, but not limited to: compliance with safety standards for light
delivery to patients such as ANSI standards, the performance capabilities of the selected
light sources, light detectors, and other devices of the system 200, light delivery rates
compatible with data acquisition rates sufficiently rapid for ime monitoring of renal
function, preserving the working life of light sources, light detectors, and other devices
related to producing and detecting light energy, and any other relevant factor.
In various s, the light sources are configured to deliver light into the
tissues of the patient 202 at a single position such as a first region 206, rated
schematically in In one aspect, the ry of light at both the excitatory
wavelength and the on wavelength to the same first region 206 enables both light
W0 2018/140984 PCT/U82018/016053
pulses to share at least a portion of the l path traveled through the s of the
patient 202 n the point of entry at the first region 206 and the point of detection at
the second region 208 and the third region 210. As discussed in detail herein below, this
arrangement of optical paths enhances the y of data produced by the system 200.
In one aspect, the first light source 218 and the second light source 220 may
be operatively coupled to a common means of light delivery. In one aspect (not illustrated)
the first light source 218 and the second light source 220 may each be operatively coupled
to a first optic fiber and a second optic fiber, respectively, and the first and second optic
fibers may be joined to a third optic fiber configured to direct light from the first optic fiber
and/or the second optic fiber into the first region 206 of the patient 202. In r aspect,
the first light source 218 and the second light source 220 may be operatively coupled to a
common optic fiber or other optical assembly configured to direct the 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 produced by the first light source 218 and the second light
source 220 may be directed in an alternating pattern into the common optic fiber or other
l assembly using an adjustable optical device including, but not limited to, dichroic
mirror or a rotating mirror.
In an aspect, the system 200 may include the sensor head 204 provided with
a sensor mount 912 configured with one or more wells within which the light sources
218/220 and light detectors 222/224 may be attached in a predetermined arrangement. In
one aspect, illustrated in and , the first light source 218 and the second light
source 220 may be situated within a source well 902 of the sensor mount 912 positioned
within the sensor head 204 (see . In an aspect, the source well 902 may contain a
first LED light source 218 producing light at the excitation wavelength and a second LED
light source 220 ing light at the on ngth operatively coupled to a single
light delivery aperture 1002 (see ) formed through the aperture plate 702, which
ensures that both wavelengths of light (i.e. tory and emission) enter the skin of the
patient 202 at approximately the same location including, but not limited to, a first region
206 as illustrated schematically in In an aspect, the source well 902 further
contains a first monitor photodiode 904 and a second monitor photodiode 906, which are
used to correct for variations in output power from the LED light sources as described in
further detail herein below.
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In an aspect, only a fraction of the light energy produced by the LED light
sources is red to the skin of the patient 202 via the single light delivery aperture
1002. In one aspect, the skin of the t 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 fraction of light produced by the LED light sources
delivered to the skin of the patient 202 may be increased by the incorporation of additional
optical elements red to focus and/or direct the light from each LED light source to
the light delivery aperture 1002. In r aspect, a diffuser may be used to mix the
output of the light s so that the light energy is rendered homogeneous at the surface
of the skin of the patient.
it) Light detectors
Referring again to the system 200 further includes a first light
detector 222 and a second light detector 224 in various s. In an aspect, the first light
or 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 comprises a 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 and emission wavelengths and
the second light detector 224 is configured to detect light received at the emission
ngth only. Combined with the illumination of the tissues of the patient 202 with
light at the excitatory wavelength only and at the emission wavelength only in an
alternating series (see the measurements from the first light detector 222 and a
second light or 224 may be analyzed as described herein below to measure the
cence of an exogenous fluorescence agent and to correct the fluorescence
measurements by ng the effects of dynamic changes in the background signal to the
correction methods described herein below.
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In various aspects, the second region 208 and third region 210 within the
tissues of the t 202, from which light is detected by the first light detector 222 and a
second light detector 224, respectively, are each separated by a nominal distance from the
first region 206 to which light produced by the first light source 218 and the second light
source 220 is delivered. This nominal separation distance may be selected to balance two
or more effects that may 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 may decrease due to light scattering along the
longer optical path between light source and light detector. This effect may be ted
by the choice of emission wavelength, which may result in a less pronounced se in
the detected fluorescence signal (i.e. light at the emission wavelength) relative to the
signals associated with detected light at the excitation wavelengths as the nominal
separation distance increases. Longer nominal separation distances result in higher
sensitivity to signal changes due to changing tissue optical properties.
In one aspect, the nominal separation distance may range from 0 mm (i.e.
colocation of light sources and light detectors) to about 10 mm. In various other aspects,
the nominal separation distance may 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 may 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 , the nominal
separation distance may be about 4 mm to balance these competing effects of logarithmic
ff of signal and reduced size of the ound signal relative to the signal from the
exogenous fluorescent agent.
ing again to the first light detector 222 may be positioned
within a first detection well 908 of the sensor mount 912 and the second light detector 224
may be positioned within a second detection well 910 of the sensor mount 912 within the
sensor head 204. The first light detector 222 and the second light detector 224 may receive
light from tissue of the patient 202 through a first detector aperture 1004 and second
detector re 1006, respectively. In an , the first detector aperture 1004, the
second detector re 1006, and the light delivery aperture 1002 are mutually separated
from one another by the l separation distance disclosed herein above including, but
not limited to, a nominal separation distance of 4 mm. In an , the first detection well
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908, second detection well 910, and light source well 902 of the sensor mount 912 may be
optically isolated from one another to ensure that light from the light s 218/220 does
not reach the light detectors 222/224 without coupling through the skin of the t 202.
The separation between the two detection wells 908/910 s that the detected
cence signal from the exogenous fluorescent agent is distinguishable from the
unfiltered excitation light, as described in detail herein below.
In an aspect, the three apertures 704 of the aperture plate 702 (see
are circular with a diameter ranging from about 0.5 mm to about 5 mm. In various other
aspects, the diameters of the apertures may 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.
In one aspect, the three apertures 704 of the aperture plate 702 are circular
apertures with a er of about 1 mm diameter. This finite width of the apertures may
result in an effective source-detector separation of less than the l separation distance
because of the logarithmic drop-off of signal with increasing separation distance from the
light sources at the skin interface of the sensor head 204.
In various aspects, the light detectors 222/224 of the system 200 may be any
suitable light detection device without limitation. Non-limiting examples of suitable light
detection devices e: photoemission detectors such as photomultiplier tubes,
phototubes, and microchannel plate detectors, photoelectric detectors such as LEDs
reverse-biased to act as photodiodes, photoresistors, photodiodes, phototransistors, and any
other suitable light detection devices. In an aspect, the light ors 222/224 are
sufficiently sensitive to detect the fluorescence emitted by the exogenous fluorescent
agents within the tissues of ts 202 that include melanin ranging from about 1% to
about 40% n in the epidermis and blood volume ranging from about 0.5% to about
2% of the skin volume. In one aspect, the light detectors 222/224 may be silicon
photomultiplier (SPM) s.
In an aspect, the first light detector 222 may be configured to detect light at
both the excitatory frequency and at the emission frequency, and the second light detector
224 may be configured to detect light at the emission frequency only. In one aspect, the
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second light detector 224 may respond only to light of the emission wavelength as a result
of the design and materials of the sensor elements of the second light detector 224. In
another aspect, the second light detector 224 may respond to a wider range of light
ngths, but may be positioned ream from an optical filter configured to pass
only the portion of incoming light with the emission wavelength and further configured to
block the e of light wavelengths e of the emission wavelength.
Any suitable optical filter may be selected for use with the second light
detector 224 to detect light ively 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, the performance of an absorption filter does
not vary significantly with the angle of incident light, whereas the performance of an
interference/dichroic filter is sensitive to the angle of incident light and may require
additional collimation optics to effectively filter the Lambertian light bution
representative of light emitted from the skin of the patient 202.
In one aspect, the second light detector 224 may be positioned downstream
of an absorptive long-pass filter configured to pass light above a predetermined wavelength
to the second light detector 224. By way of non-limiting example, the second light or
224 may be positioned downstream of an long-pass OG530 filter configured to pass light
with wavelengths above about 530 nm. Other non-limiting examples of suitable filters
include a Hoya 054 filter and a Hoya CM500 filter.
In s aspects, an l filter 244 configured to absorb excitation
wavelength light may be positioned within the second detection well 910 between the
second light detector 224 and the second detector aperture 1006. In one aspect, the optical
filter 244 may be constructed from OG530 Schott glass. The thickness of the l filter
244 may be selected to enable an optical density ent to filter the excitation light by
about three orders of magnitude. In one aspect, the thickness of the optical filter 244 may
range from about 1 mm to about 10 mm. In various other aspects, the ess of the
optical filter 244 may range from about 1 mm to about 8 mm, from about 2 mm to about 6
mm, and from about 3 mm to about 5 mm. In various additional aspects, the thickness of
the optical filter 244 may be about 1 mm, about 2 mm, about 3 mm, about 4 mm, about 5
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mm, about 6 mm, about 7 mm, about 8 mm, about 9 mm, and about 10 mm. In one aspect,
the optical filter 244 is a 3-mm thick filter constructed of OG530 Schott glass.
In an additional aspect, an optical diffuser may be provided within the light
source well 902. In this aspect, the optical diffuser enables mixing of the light entering the
light source well 902 from the first and second light sources 218/220. By mixing the light
from the first and second light sources 218/220 using the optical diffuser prior to
illumination of the first region 206 of the patient 202, the similarity of the optical paths
taken by emission-wavelength light and tion-wavelength light through the tissues of
the patient is enhanced relative to the corresponding l paths taken by unmixed light,
thereby reducing a potential source of variation.
In an aspect, a arent material configured to pass light of both the
excitatory and emission wavelengths may be positioned within the first detection well 908
between the first light detector 222 and the first detector aperture 1004. In this , the
transparent al may be any material with similar l properties to the material of
the l filter 244 including, but not limited to, thickness and index of refraction. In one
aspect, the arent al within the first detection well 908 may be fused silica glass
of the same thickness as the optical filter 244.
By way of non-limiting example, the transmission spectrum of the 0G 530
filter is provided in As illustrated in the transmission spectrum of the 0G
530 filter overlaps with the emission spectrum of the 1Vfl3-102 exogenous fluorescent agent
and the emission spectrum of a green LED used as a second light source 220 (emission
wavelength). In addition, the transmission spectrum of the 0G 530 filter excludes the
emission spectrum of the blue LED used as a first light source 218 and the absorbance
spectrum of the MB-102 exogenous fluorescent agent (excitation wavelength).
In an aspect, the transparent material such as glass 246 and the optical filter
244 may be secured to ledges formed within the first detection well 908 and the second
detection well 910, respectively. The transparent material such as glass 246 and the optical
filter 244 may 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 the first
detector aperture 1004 and the second detector aperture 1006 travels through the optical
filter 244 or glass 246 before detection by the first and second light ors 222/224. In
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another aspect, the sides of the optical filter 244 or glass 246 may be painted black with a
light-absorbing coating including, but not limited to, India ink to ensure that light does not
reach the first and second light detectors 222/224 without passing through the optical filter
244 or glass 246.
In an aspect, the height of the detection wells 908/910, combined with the
er of the detector apertures 1004/1006 may limit the fraction of the light emitted
from the second region 208 and third region 210 of the patient’s skin that reaches the active
areas of the light detectors 222/224 due to the Lambertian distribution of the angle of the
light leaving the patient’s skin. In one aspect, the on of light emitted from the second
region 208 and third region 210 of the patient’s skin received by the light detectors 222/224
may range from about 5% to about 90%. In various other aspects, the fraction of light may
range from about 5% to about 15%, from about 10% to about 20%, from about 15% to
about 25%, from about 20% to about 30%, from about 25% to about 35%, from about 30%
to about 40%, from about 35% to about 45%, from about 40% to about 60%, from about
50% to about 70%, and from about 60% to about 90%.
In one aspect, for the sensor head 204 rated in and with
1-mm diameter apertures 1002/1004/1006, about 10% of the light emitted from the surface
of the patient’s skin may reach the active area of the light ors 222/224 to be detected.
In various aspects, the sensor head 204 may r include additional l elements
including, but not limited to, lenses and/or prisms configured to compensate for the
Lambertian distribution of light angles in order to enhance the on of light emitted
from the patient’s skin that is directed to the active area of the light detectors 4.
iii) Temperature sensors
Referring to the sensor head 204 may r include one or more
additional temperature sensors 228 configured to monitor temperatures of various regions
within the sensor head 204 and in the vicinity of the sensor head 204. Non-limiting
examples of suitable s for which the temperature may be monitored by the one or
more additional temperature sensors 228 include: temperature at the skin surface of the
patient 202, temperature in the vicinity of the first light source 218 and/or second light
source 220, ambient temperature outside of the sensor head 204, temperature of housing
600 of sensor head 204, and any other suitable region. In one aspect, additional
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temperature sensors 228 may be configured to r the temperatures in the vicinity of
temperature-sensitive electrical components including, but not limited to: light sources
218/220 such as LEDs, 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 one or more additional temperature sensors
228 may be used as feedbacks in a control method for one or more of the temperature-
sensitive s of the system 200 as described herein below.
By way of non-limiting example, a temperature measurement may be used
to control the amount of light energy produced by an LED used as a first or second light
source 218/220. In this example, LED temperatures measured by an second temperature
sensor 1108 (see ) may be used in a control scheme to modulate the amount of
power ed to an LED light source to compensate for the effect of LED temperature on
the light output of the LED. In another aspect, additional ature sensors 228 may
monitor the temperatures of LED light sources 218/220 to monitor and/or compensate for
temperature variations of the LEDs as well as to monitor and/or sate for
temperature-dependent transmission of the optical filters to maintain relatively constant
output wavelengths.
By way of another miting example, an additional ature sensor
228 may be included in the sensor head 204 in the form of a temperature sensor 816 (see
configured to monitor the temperature of the housing 600 in the vicinity of the
t surface 606 of the sensor head 204. ing to FIG, 8, and the
ature sensor 816 may be epoxied into the temperature sensor opening 706 in the
aperture plate 702 in one aspect. In this aspect, the space 918 between the circuit board
(not shown) and the lower housing 604 may be filled with a thermally conductive putty to
ensure good thermal conduction and dissipation.
In this example, the measured housing ature may 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 r aspect, additional temperature sensors 228 may monitor the
temperatures of LED light sources 218/220 to r and/or compensate for temperature
variations of the LEDs to enable the maintenance of relatively constant output wavelengths
by the LED light sources 218/220.
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In an additional aspect, temperatures measured by one or more additional
temperature sensors 228 may provide for subject safety by disabling one or more electrical
devices ing the light sources 218/220 and/or light ors 4 if an over-
temperature condition is detected. In one aspect, an over-temperature condition may be
indicated if the housing temperature detected by the temperature sensor 816 is greater than
about 40°C. In various other aspects, an over-temperature condition may be detected of the
housing ature is greater than about 405°C or greater than about 41 .00 C.
B. Controller
Referring again to the system 200 in s aspects may e a
controller 212 configured to operate the light sources 218/200 and light detectors 222/224
in a coordinated fashion to obtain a plurality of measurements used to obtain the
fluorescence of the exogenous fluorescent agent within the tissues of the patient 202, to
correct the fluorescence data to remove the effects of dynamic changes in the background
signal as described herein below, and to transform the fluorescence ements into a
parameter representative of the renal function of the patient 202. is a schematic
diagram of an electronic circuit 1100 that rates the arrangement of various electrical
ents that enable the operation of the system 200 in an aspect. In one aspect, the
ller 212 may be a computing device further including an operation unit 214 and a
y unit 216.
1) Light source control unit
Referring again to the controller 212 may include a light source
control unit 230 configured to operate the first light source 218 and the second light source
220 to produce light at the excitation wavelength and emission wavelength, respectively in
a coordinated manner to produce a repeating pulse sequence as illustrated schematically in
In various aspects, the light source control unit 230 may produce a plurality of
light control signals encoding one or more light control ters including, but not
limited to: activation or deactivation of each light , 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 parameter associated with light pulse fluence or
light pulse power, other light source-speciflc parameters controlling the light output of the
light source, and any other relevant light control parameter. In an aspect, the light source
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l unit 230 may receive one or more feedback measurements used to modulate the
ity of control signals to compensate for variations in performance of the light sources
in order to maintain a relatively stable output of light from the light sources. Non-limiting
examples of feedback measurements used by the light source control unit 230 include: light
output of the light sources 0 measured within the source well 902 by the first
monitor photodiode 904 and the second monitor photodiode 906, respectively,
temperatures of the light sources 218/220, and any other feedback measurement relevant to
monitoring the mance of light sources 218/220.
By way of non-limiting example, the light source control unit 230 may be
configured to e LED light sources 218/220. In this e, the light output of the
LED light sources 218/220 may be controlled by lling the magnitude of current
provided to each LED. In an aspect, the light source control unit 230 may include at least
one waveform generator 1122 including, but not limited to, a field programmable gate
array FPGA with a 16-bit DAC 1124 operatively coupled to a LED current source 1126, as
illustrated in . In an , waveforms generated by the at least one waveform
generator 1122 including, but not limited to square waves, may control the output from the
LED current source 1126. In an aspect, the magnitude of the current supplied to the LED
light sources 218/220 may be adjustable based on the rm signals ed by the
waveform generator/FPGA 1122.
Referring to in one , each light pulse sequence 500 includes an
emission wavelength light pulse 502 and an excitatory wavelength light pulse 504 that are
both made up of a plurality of square waves 506 produced by the first and second LED
light sources 218/220. Referring to , square waves generated by the waveform
generator 1122 are ed by the LED current source 1126. The current generated by the
LED current source includes a square waveform similar to the waveform generated by the
waveform generator 1122. Without being limited to any particular theory, because the
intensity of light produced by the LED light sources 218/220 is proportional to the
magnitude of the current received, the light ed by the LED light s 218/220
also includes the square waveform as illustrated in In another aspect, discussed in
additional detail below, the square waves produced by the rm generator 1122 may
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, the detection of
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ambient light, from the detector signals generated by the light detectors 222/224 during
illumination of the tissues of the patient at the emission and tory wavelengths by the
first and second light sources 218/220, respectively.
In various other aspects, a variety of alternate LED pulse modulation
schemes may be equivalently employed without limitation. In one aspect, the excitation
and emission pulses are delivered in an alternating series persed with a dark period
after each pulse. In another aspect, the first and second LED light sources 218/220 are
each modulated with a 50% duty cycle but at different tion frequencies, allowing
the signals associated with the excitation and emission pulses to be separated by frequency
filtering.
Without being limited to any particular theory, the overall optical power
delivered to the patient’s skin may be d by at least two factors: photobleaching of the
exogenous fluorescent agent and/or endogenous phores, as well as overheating of
the patient’s tissues nated by the system 200. In one , tissue heating may
impose an absolute limit of about 9 mW on the optical power that can be delivered to the
skin, based on safety standards including, but not limited to, ANSI/IESNA RP-27.1-05. In
another , leaching of the skin autofiuorescence associated with endogenous
chromophores including, but not d to, collagen, obin, and melanin may
contribute a background signal to the measured fluorescence that remains vely
constant so long as no eaching of the chromophores occurs. This constant
autofiuorescence background may be subtracted from the raw fluorescence signal, but if
autofiuorescence varies over time due to photobleaching, this background correction may
interfere with the c calculation of the renal decay time constant (RDTC). In an aspect,
the light output power of the first light source 218 and/or second light source 220 may be
limited to levels below power thresholds associated with chromophore photobleaching.
Referring again to the light output of the light sources 218/220 may
be measured using monitor photodiodes 904/906 in various aspects. Because the light
ity reaching these monitor photodiodes 904/906 is typically much stronger than the
light intensity that reaches the light detectors 222/224 through the patient’s skin, less
sensitive light detecting devices including, but not limited to, PIN photodiodes may be used
to monitor the output of the light sources 218/220.
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In various aspects, the system 200 may be configured to operate over a
range of skin tones observed in the human population. Without being limited to any
particular theory, variations in skin tones between ent patients 202 may result in
ions in the detected fluorescence signals ranging over about three orders of
ude. In addition, variations in the concentrations of exogenous fluorescent agent
within each patient 202 may vary over a range of about two orders of magnitude due to
renal elimination of the agent over time. In various aspects, the system 200 may be
configured to detect cence from the endogenous fluorescent agent over an intensity
range of more than five orders of magnitude. In these various aspects, the system 200 may
be configured by modulation of at least one operational parameter including, but not
limited to: magnitude of light output by the light sources 218/220 and sensitivity of light
detectors 222/224 ponding to detector gains.
In one aspect, the intensity of the light output by the light sources 218/220
may be manually set by a user via the operation unit 214. In another aspect, the light
source control unit 230 may be configured to te the intensity of light produced by
the light sources 218/220 automatically. In an aspect, the light source control unit 230 may
be configured to control the light intensity produced by the LED light sources 218/220
within a range of normalized output intensities from 0 (off) to 1 (maximum . In an
, the ity of the light sources 218/220 may be set by the light source control unit
230 in coordination with the detector gains of the light detectors 222/224 set by the light
detector control unit 232, as described herein below.
In one aspect, signals ed during the first 10 detection cycles ed
by the system 200 after initialization of data ition, but prior to the injection of the
exogenous fluorescent agent, may be used by the light source l unit 230 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 cycle may be
obtained with the LED light sources 218/220 set at about 10% of maximum LED intensity
(corresponding to a normalized output intensity of 0.1) and with a low gain setting for the
light detectors 222/224. Based on the detected intensity of light received at the light
detectors 222/224 at the excitation and emission wavelengths for one ion cycle, the
corresponding LED intensities may be modulated to enable the analog signals produced by
the light detectors 222/224 to correspond to about 1A of the full range of each detector
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analog-to-digital convertor (ADC) at the low detector gain setting. If the signals produced
by the light detectors 222/224 in se to the light produced by the second LED light
source 220 at the emission wavelength do not agree, the larger signal may be used to
modulate the power setting of the second LED light source 220. If the method described
above results in modulation to an LED intensity g higher than the maximum intensity
(corresponding to a normalized output intensity of 0.1), the LED intensity setting is set to
the maXimum setting. t being limited to any particular theory, the targeted levels of
signals produced by the light detectors 222/224 (i.e. 1A of the ADC range) is selected to
reserve additional light detection capacity to detect signals resulting from variations in
optical properties of the tissues of the patient 202 during the study due to any one or more
of a plurality of factors including, but not limited to, the introduction of the exogenous
fluorescent agent into the patient 202.
In the above one aspect, once the LED intensities are set by the light source
l unit 230 in coordination with the detector gains of the light detectors 222/224 set by
the light detector l unit 232 over the first 10 ion cycles, an additional 10
ion 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, followed by a
recalculation of the LED intensity settings and detector gains as described . If the
newly calculated LED intensity is within a factor of two of the previously determined
setting, and the detector gains are not d, the previously determined settings are
maintained for subsequent data ition cycles used to determine renal function.
Otherwise, the settings are updated using the same method described herein and another 10
data acquisition cycles conducted to confirm the ity of the settings. This process
repeats until either the settings are determined to be acceptably stable or 10 data acquisition
cycles are conducted to obtain the settings, in which case the most recently determined
settings are used for all subsequent data acquisitions, and the user may be notified via the
display unit 216 that the gs may not be optimal.
it) Light detector control unit
Referring again to the controller 212 may include a light detector
control unit 232 configured to operate the first light detector 222 and the second light
or 224 to enable the detection of light at the emission wavelength and unfiltered light
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at all wavelengths, respectively. In various s, the light detector control unit 232 may
produce a plurality of detector control signals encoding one or more detector control
parameters including, but not limited to, detector gains. In various other aspects, the light
detector control unit 232 may produce a plurality of light measurement signals encoding
the ity of light detected by the light detectors 222/224 including, but not limited to
raw detector signals that may be received by an analog-to-digital convertor (ADC) 1102
(see ) in various aspects. In another aspect, the detector gains and/or other detector
control signals may be manually set by a user or gains when the system 200 is
configured in an Engineering Mode.
In various other aspects, the amount of light received by the light detectors
222/224 may vary due to any one or more of at least l s including, but not
limited to: variation in skin tones observed between individual patients 202, variations in
the concentrations of exogenous fluorescent agent within each patient 202, and any other
relevant parameter. In one aspect, gains of the first light detector 222 and the second light
detector 224 may be set by a user via the operation unit 214. In r aspect, the light
detector l unit 232 may be configured to modulate the gain of the light detectors
222/224 automatically via a bias voltage gain of the bias voltage generator 1112 (see ).
In one aspect, signals obtained during the first 10 detection cycles ed
by the system 200 after initialization of data acquisition, but prior to the injection of the
exogenous fluorescent agent, may be used by the light detector control unit 232 to
automatically adjust the gains of the light detectors 222/224, as well as the output
intensities of the light s 218/220. As described herein previously, the l detection
cycle may be obtained with the LED light sources 0 set at about 10% of maximum
LED ity sponding to a normalized output intensity of 0.1) and with a low gain
setting for the light detectors 222/224 and the LED intensities may be modulated to enable
the analog signals produced by the light detectors 222/224 to correspond to about 1A: of the
full range of each detector analog-to-digital convertor (ADC) at the low detector gain
setting.
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In this one aspect, if the intensity of the first LED light source 218
cing light at the excitation ngth) is set to the maximum of the LED power
range, a high detector gain may be considered for the second light detector 224
corresponding to the filtered measurements of the excitation wavelength only. In various
aspects, the high detector gain may be 10-fold higher than the corresponding low detector
gain for a given light detector. Without being limited to any particular theory, the expected
peak ed fluorescence signal from the exogenous fluorescence agent over the course
of injection and renal ation is typically expected to be about 10% of the magnitude of
the signal received during illumination at the excitation wavelength by the first light source
218, assuming that the exogenous fluorescence agent is MB-102 introduced into the patient
202 at a dose level of about 4 umol/kg of patient weight. In an aspect, if the expected
detector signal received during illumination at maximum LED intensity and with the
detector gain set to the high setting remains below 10% of the range of the detector ADC,
the detector gain for that measurement be increased by ten-fold. In another aspect, the
saturation ion may persist for a pre-defined period of time ing, but not limited
to, a 30-second period before adjustments are made to the detector gain or LED power to
avoid reacting to spurious signal spikes.
In another , the light detector control unit 232 may adjust the detector
gain to a lower gain level if the ed light signals from one of the light detectors
4 exceed a threshold percentage of the maximum ADC range to avoid signal
saturation. Although the highest threshold percentage of the maximum ADC range
associated with signal tion is 100%, the onset of severe detector non-linearity takes
place at threshold percentages of about 40% or more, and mild detector non-linearity
occurs at threshold percentages in excess of about 15%. In various aspects, the threshold
percentage of the maximum ADC range may 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 ed light signals from one of the light detectors 222/224
exceed about 8% of the maximum ADC range, the gain setting will be adjusted. By way of
non-limiting example, if the or gain on the nearly saturated signal is high, it will be
adjusted to low. If the current detector gain is set to low and the corresponding detected
light signal remains above the threshold percentage of the maximum ADC range, the LED
output power setting of the corresponding LED light source may be reduced ten-fold.
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In an aspect, the light or control unit 232 may receive one or more
feedback measurements used to modulate the ity of detector signals to sate
for variations in the mance of the light detectors due to variations in temperature
and/or light source output. Non-limiting examples of feedback measurements used by the
light detector control unit 232 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 (see ), temperatures of the light detectors 222/224
measured by a first temperature sensor 1106, LED temperatures measured by a second
temperature sensor 1108, temperature of the sensor head housing measured by a third
temperature sensor 1128, LED supply current from the LED current source 1126, and any
other feedback measurement nt to monitoring the performance of light detectors
222/224.
In various aspects, the light detectors 222/224 may be n photon
multiplier (SPM) detectors that may e low-noise internal amplification, and may
function at lower light levels relative to other light sensor devices such as PIN photodiodes.
The detector signal generated by the SPM detectors 222/224 may be amplified using
mpedance amplif1ers 1120/1118, respectively (see ) to ate a t
generated by each SPM light detector 222/224 into a measurable detector voltage. The
transimpedance amplif1er 1118 on the second SPM light detector 224 (i.e. detects filtered
lights at the excitation wavelength only) may include a switchable detector gain that may
select a low gain configured to detect a larger dynamic range for fluorescence
measurements when the first LED light source 218 is activated to e light at the
emission wavelength. The switchable detector gain that may further select a high gain
setting for the second SPM light detector 224 when the second light source 220 is inactive
to enhance the sensitivity of the second SPM light detector 224 during the phase of the
detection cycle when light at the emission wavelength produced by the exogenous
fluorescent agent within the tissues of the patient 202 is detected, to ensure that the
expected dark current from the second SPM light detector 224 es less than 1A: of the
total ADC output range. In one aspect, the second transimpedance amplifier of the second
SPM light detector 224 may include a low detector gain configured to provide a
transimpedance gain of about 4 k9 ponding to about twice the value of the
transimpedance resistor due to differential operation, and may further include a high
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detector gain configured to e a transimpedance gain of about 40 k9. In another
aspect, the first mpedance amplifier of the first SPM light detector 222 may include a
fixed detector gain configured to provide a transimpedance gain of about 2 k9.
iii) Acquisition Unit
Referring again to the controller 212 may further include an
acquisition unit 234 in various aspects. The acquisition unit 234 may be configured to
receive a plurality of signals from the light sources 218/220, light detectors 222/224, and
additional light detectors 226 and onal temperature sensors 228 and processing the
plurality of s to produce one or more raw signals including, but not limited to, raw
fluorescence signals encoding the intensity of fluorescence detected by the second light
or 224 during nation at the excitation wavelength, and raw internal reflectance
signals corresponding to the intensity of light at the excitation wavelength ed by the
first light detector 222 during illumination at the excitation wavelength as well as the
intensity of light at the emission wavelength detected by the both light detectors 222/224
during illumination at the emission wavelength.
The plurality of signals received from the various sensors and s
bed herein above are typically analog signals including, but not limited to, electrical
voltages and currents. In various aspects, the acquisition unit 234 may enable the
transmission of the analog signals to one or more analog-to-digital converters (ADCs) to
convert the analog signals into digital signals for subsequent processing by the processing
unit 236. is a schematic diagram of a circuit 1100 illustrating the arrangement of
various electrical devices and components of the sensor head 204. In one aspect, the
analog signals encoding the intensity of light detected by the first light detector 222 and the
second light detector 224 may be received by a first ADC 1102.
In various aspects, the analog signals produced by the light detectors
222/224 and various monitor s may be digitized using at least one 24-bit Sigma-
Delta ADC. Referring again to , analog s encoding the measurements from
time-sensitive sensors may be digitized using a high-speed 24-bit Delta ADC 1102
in one aspect. In this aspect, time-sensitive sensors include sensors associated with the
production and detection of light pulses characterized by potentially y-changing
signals. Non-limiting examples of time-sensitive sensors of the system 200 include: first
W0 40984
and second light detectors 1118/1120, and first and second monitor photodiodes 904/906.
In another aspect, analog signals encoding the ements from less time-sensitive
sensors may be digitized using a low-speed 24-bit Sigma-Delta ADC 1104. In this other
aspect, the less time-sensitive sensors include sensors associated with monitoring system
conditions characterized by lly slow-changing signals including, but not d to,
temperatures of various system ents and/or regions. Non-limiting examples of less
time-sensitive sensors of the system 200 include: a first and second temperature sensor
108 configured to monitor the temperatures of the light detectors 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.
In various aspects, the ition unit 234 may be further configured to
enable synchronous detection of light by detectors 222/224. Without being limited to any
particular theory, synchronous detection methods are t to reject noise from the
or signals associated with the detection of light produced by the light sources
118/120 and fluorescence produced by the exogenous fiuorescent agents within the tissues
of the patient 202 by distinguishing the detector signals from noise associated with the
detection of ambient light or other sources of interference.
is a schematic illustration of a synchronous detection method in one
aspect. Referring to and , the waveform generator/FPA 1122 may generate
a digital square wave 1202 that is received by the DAC 1124, and the resulting analog-
converted square wave is received by the LED current source 1126. The ing current
produced by the LED current source 1126, also terized by a waveform proportional
to the analog-converted square wave drives LED light s 0. The light produced
by LED light sources 218/220, after passing through the tissues of the patient 202 are
detected, along with the fluorescence produced by the endogenous fiuorescent agent, by the
light detectors 222/224 and are digitized by the high-speed ADC 1102.
Referring again to and , the digital square wave 1202
generated by the waveform generator/FPA 1122 may also be converted by a DAC 1110
(see ) to an in-phase reference sine wave 1210 and an out-of-phase/quadrature
reference cosine wave 1212. In an aspect, the digitized detector signals from the ADC
1102 and the in-phase reference sine wave 1210 may be sampled and subjected to signed
W0 2018/140984
multiplication at a first multiplier 1214 to generate a plurality of in-phase modulated
signals. In addition, the digitized detector signals and the quadrature reference cosine wave
1212 may be sampled and subjected to signed multiplication at a second multiplier 1216 to
generate a plurality of quadrature (out-of-phase) modulated signals. In this aspect, the
acquisition unit 234 may delay the samples from the reference waves 1210/1214 by an
amount equivalent to the relative delay between the DAC 1124 generating the reference
waves 1210/1214 and the ADC 1102 digitizing the detector signals to synchronize the
reference waves 1210/1214 to the detector data being acquired.
Referring again to , the in-phase modulated signals may be summed
in a first accumulator 1218 to generate an se intensity signal 1224. Similarly, the
quadrature modulated signals may be summed in a third lator 1222 to generate a
quadrature intensity signal 1228. The raw digitized detector signal may also be summed in
a second lator 1220 to generate an average intensity signal 1226. In on, the
in-phase intensity signal 1224 and the quadrature intensity signal 1228 may be root-sum
squared to generate a magnitude signal 1230.
t being limited to any particular theory, the integration interval of the
accumulators 1218/1220/1222 may correspond to an integer number of modulation cycles
sponding to cycles of the l square wave 1202) to avoid a bias on the measured
signal. The phase accumulators 220/1222 used to control the synchronous detection
operates on integer numbers, but the sample clock frequency and the modulation frequency
are not integer-divisible, so the number of cycles is not exactly an integer. However, the
error associated with this mismatch may be minimized by adjusting the actual modulation
ncy to match as y as possible with the achievable sampling als and
allocating an appropriate number of bits to the phase accumulator. In one aspect, the error
associated with the mismatch between the modulation frequency and the sampling intervals
may be on the order of about one part in 106.
In one aspect, the digital square wave 1202 used to modulate the LED light
sources 218/220 and to enable onous detection method as described herein above is
ed at a frequency of about 1 kHz. t being limited to any particular theory, a
square wave was selected as the modulating waveform to enable an enhancement in signal
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to noise ratio (SNR), as compared to a pure sinusoidal wave as the modulating waveform
for the same peak power level.
In another aspect, the acquisition unit 234 may be further configured to
enable demodulation of the in-phase intensity signal 1224, average ity signal 1226,
and quadrature intensity signal 1228. In one aspect, the acquisition unit 234 may pick out
each component at the fundamental harmonic, which is characterized by an amplitude that
is (4/71) times larger than the amplitude of the square wave 1202 used to te the
intensity signals 1224/1226/1228. In various aspects, to reject 50/60 Hz electrical noise
generated by the alternating current electrical power sources, and corresponding 100/120
Hz optical noise generated by ambient light sources powered from those electrical power
sources, the integration period of the accumulators 1218/1220/1222 may be selected to be a
multiple of 100 ms. In these various aspects, this selected integration period ensures that
integration by the accumulators 1218/1220/1222 occurs over an integer number of cycles
for the 50, 60, 100, and 120 Hz signals.
iv) sing unit
Referring again to the controller 212 may further
include a processing unit 236 configured to apply tions to the demodulated
detector signals and to transform a ed portion of the ted detector signals into
a measure of renal function in various aspects. is a block diagram illustrating
the subunits of the processing unit 236 in an aspect. Referring to , the
processing unit 236 may include a pre-processing subunit 1302 configured to determine
and correct the detector s to remove signal artifacts associated with a variety of
confounding effects including, but not d to, physiologically-induced signal
variations, variations in power supplied to the light sources 218/220, nearities in
detector se, ambient temperature variation, and tissue heterogeneity. The
processing unit 236 may further include a background subtraction subunit 1304
configured to remove the portion of the detector signals attributable to background
factors such as autofiuorescence of the tissues and/or leakage of light at the excitation
wavelength through the optical filter 244 of the second light detector 224. The
sing unit 236 may additionally include a background correction t 1306
configured to enable a method of applying a background correction method to remove
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the effects of dynamic s in the background signal related to changes in
autofiuorescence and/or the leak-through of tory-wavelength light to the second
light detector 224 configured to detect emission-wavelength light only, and to apply the
background correction to the first detector, turning DRemm into DRewhotm. The
sing unit 236 may further include a post-agent administration selection subunit
1308 configured to select a portion of the detector data associated with the post-
equilibration period for subsequent analysis to ine renal on of the patient.
The processing unit 236 may further include an RDTC calculation subunit 1310
configured to transform the detector signals obtained over the post-equilibration period
to e a renal decay time constant indicative of the renal function of the patient.
The processing unit 236 may also include a fault detection subunit 1312 configured to
monitor the magnitudes of the detector signals to detect any malfunctions of the system.
- pre-processing subunit
In one aspect, the raw signals corresponding to the light intensity ed
by light detectors 222/224 ponding to illumination by the first light source 218 and
the second light source 220 at the tion and emission wavelength, 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, resulting in signals that
more accurately reflect the underlying specific signals of interest.
By way of several non-limiting es, the intensity of light produced by
a light source may vary due to one or more of a plurality of factors including, but not
limited to: fluctuations in the electrical current supplied to the light source and variations in
the ambient temperature of the light source. Light characterized by two or more
wavelengths emanating from the same source re of the sensor head may not share the
same path to the same detector. The detectors may have thermally-dependent sensitivity
and gain. Further, the l filter associated with the second light detector 224 may have
ature-dependent transmission properties.
In one aspect, the pre-processing subunit 1302 is configured to process the
raw signals corresponding to light intensities detected by the first and second light
detectors 4 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
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not limited to, the plurality of factors described above. A is a block m
illustrating the modules of the pre-processing t 1302 in one aspect. B is a
block diagram illustrating the modules of the pre-processing subunit 1302a in a second
aspect.
In one aspect, illustrated in A, the pre-processing subunit 1302 1)
resamples the signals using the methods of the resampling module 2202 as described
below, 2) removes saturated or signals using the methods of the detector output
saturation detection and removal module 2204 as described below, 3) corrects for
temperature-dependent detector gain using the methods of the detector temperature
correction module 2206 described below, 4) corrects the signals for instrument light
ionality using the methods of the light directionality correction module 2208
described below, 5) corrects the signals for filter throughput and temperature-dependent
variation of fluorescence light using the methods of the filter throughput temperature
correction (emission) module 2212 described below, 6) corrects for tissue geneity
using the methods of the tissue heterogeneity tion module 2216 described below, 7)
corrects the signals for filter throughput and temperature-dependent variation of excitation
light and signal decomposition using the methods of the filter throughput temperature
correction (excitation) module and signal decomposition module 2214 as described below,
8) corrects for optical power variation using the methods of the fractional photon
normalization module 2218 as described below.
In one aspect, rated in B, the ocessing subunit 1302a
calculates signal magnitudes using the methods of the detector temperature correction
module 2206a as described below, resamples the signals using the methods of the
resampling module 2202a as described below, removes saturated samples using the
methods of the detector output saturation ion and removal module 2204a as described
below, corrects the signals for temperature-dependent or gain using the methods of
the or temperature correction module 2206a described below, corrects the signals for
optical power variation using the methods of the fractional photon normalization module
2218a as described below, corrects for excitation light leakthrough onto the measured
cence signal using the filter hput temperature correction ation) module
and signal decomposition module 2214a as described below, and corrects for fluorescence
W0 2018/140984 PCT/U82018/016053
light leakthrough onto the measured tion diffuse reflectance signal using the filter
throughput temperature correction (emission) module 2212a as described below.
- resampling module
Referring to A and B 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 t 202 ing, but not
limited to, heartbeat and breathing. Typically, an acquisition sequence is characterized by
alternating interval of illumination at the excitation and emission separated by intervals of
no illumination (i.e. dark intervals). gh both illumination intervals
ation/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 n the tion and emission illumination intervals. t
being limited to any particular theory, if the separation interval ated with an
ition sequence is on the order of a separation interval between physiological events,
such as heartbeats or respiration, logical noise may be introduced to the s. In
various s, this physiological noise may be reduced by resampling the signals
associated with the excitation and emission illumination to overlap prior to subsequent
processing of the signals.
By way of non-limiting example, a sample sequence may include a 100 ms
dark interval, a 100 ms interval of illumination at the excitatory wavelength, a second 100
ms dark interval, and a 100 ms interval of illumination at the emission wavelength. Each
sample packet is logged with a single timestamp, and each sample packet is separated by a
400 ms interval. Because physiological signal variations, such as from heartbeats, occur on
this same timescale, the 200 ms difference between signal ition associated with the
excitatory and emission wavelengths becomes apparent in the signals. This physiological
signal noise may be reduced using the pre-processing subunit 1302 by first resampling the
signals ated with illumination at the excitatory and emission wavelength illumination
to overlap prior to performing any additional signal processing as described below. In this
non-limiting example, the signals associated with illumination at the excitatory wavelength
may be shifted forward by 100 ms and the signals associated with illumination at the
W0 2018/140984 PCT/U82018/016053
emission wavelength may be shifted backwards by 100 ms, resulting in an overlap of the
signals.
In various aspects, the resampling module 2202 performs resampling as
described above on s 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.
- detector output saturation detection and removal module
Referring again to A and B the pre-processing subunit
1302/1302a in s aspects es 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 detectors 222/224. In one aspect, the pre-processing t
1302 compares the detected signals to the m 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 and removes that
value from further processing.
- detector temperature correction module
Referring again to A and B the pre-processing subunit
302a in various aspects includes a detector temperature correction module
2206/2206a configured to enable a temperature correction to compensate for the thermal
sensitivity of the light detectors 222/224. In one aspect, the intrinsic detector gain for a
silicon photomultiplier (SPM) device typically used as a light detector is proportional to the
difference between the device own voltage and the bias e applied by the bias
voltage generator 1112 (see ), ed to herein as an overvoltage. In this aspect,
the breakdown voltage varies with ature in a haracterized manner. In one
, the temperature tion accounts for both this internal detector gain variation and
additionally temperature-related variation in the photon detection efficiency.
In one aspect, the temperature correction may be a scaling correction
applied to the detector measurements in which the scaling correction is based on a
measured detector temperature. In an aspect, the measured light detector signals may be
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divided by the calculated gain C(Z) to remove the temperature dependency. The scaling
correction C(Z) may be ated according to Eqn. (2):
C(71) : Cv' Vbias _ Vbreakdown(1 + CT)T—TO Eqn~ (2)
In Eqn. (2), the monitor temperature T is obtained from a first temperature
sensor 1106 (see ) configured to monitor the temperature of the sensors 4.
The bias e (mes) may be measured by the bias voltage generator 1112. The
breakdown voltage (mekdown) and reference temperature (T0) are constants specific to the
particular light detector device included in the system 200. By way of non-limiting
example, if the light detectors 4 are silicon photomultiplier (SPM) devices, Vbreakdow”
may be 24.5 V and T0 may be 21 degrees C. In another aspect, the coefficients CV and CT
used in Eqn. (2) may be derived cally based on measurements obtained using a
constant phantom over an ambient temperature ranging from about 18 degrees C to about
26 degrees C.
In another aspect, the temperature portion of the gain correction is
determined by the Eqns. (3)—(5).
GuseCase — Cv_ T —T
' Vbiasmeaswed _ Vbreakdown(1 + CT) measured 0 Eqn~ (3)
Gnominal — Cv_ T - —T
' ominal _ down(1 + CT) nominal 0 Eqn~ (4)
GuseCase
G _
correctlon — ECln~ (5 )
Gnominal,
This gain correction can be applied to each of the signal magnitudes as
measured by the first and second light detectors 222/224 as follows:
SPMmagnitude
5PMmagnitudecorrected = Eqn. (6)
Gcorrection
In an aspect, the magnitudes of the ements from each or and
monitor photodiode are calculated from the root sum-squares of the in-phase magnitude
signals 1230 (I) and quadrature magnitude signals 1232 (Q) according to Eqn. (1):
M = I2 + Q2 Eqn. (1)
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The signal magnitudes from the light detectors 222/224 calculated using
Eqn. (1) are normalized by the monitor photodiode ude for each measurement set
corresponding to the measurements obtained during illumination by one of the LED light
sources 218/220 at either the tion or emission wavelength. Because both monitor
photodiodes 904/906 may positioned in the same source well 902 as both LED light
sources 218/220 (see , the average of the two monitor photodiode udes from
the corresponding measurement set is used.
In an aspect, the in-phase intensity signal 1224, quadrature intensity signal
1228, and average intensity signal 1226 (see ) are further processed for the number
of accumulated samples and ADC scaling such that the ity signals 1224/1226/1228
are ed as fraction of the full range of the peed ADC 1102 (i.e. ranging from a
minimum of 0 to a maximum of 1). The measurements of the monitor photodiodes
904/906 (see ) are similarly scaled as a fraction of the full range of the low-speed
ADC 1104.
In one aspect, Gcorrection may incorporate a power correction to t for
the effects of fluctuations in the LED power supply. In this aspect, the signals from the
first monitor photodiode 904 and the second monitor photodiode 906 are calibrated by
ing optical output power with a power meter as light intensities from the light
sources 218/220 are varied. The calibration coefficients for each light source 218/220,
Cmm] and Cmmg, are calculated as detector-measured milliWatts per recorded monitor
iode signal value. Cmm] and Cmmg are used to determine the absolute light output
into tissue at each wavelength.
Referring again to B, the detector temperature correction module
2206a corrects signal magnitudes for the varying ity of the LEDs by normalizing the
temperature-corrected detected s using the LED output signal PDmagm-tude measured
by the first monitor photodiode 904 and/or the second monitor photodiode 906. In this case,
the Gammon variable for each light source 218/220 from above is amended as follows:
GuseCase
Gcorrection —_ * PDmagnitude Eqn~ (7)
Gnominal
- light directionality correction module
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Referring again to A, the pre-processing subunit 1302 in this aspect
includes a light directionality correction module 2208 configured to enable a tion to
variations in the detected signals associated with differences in the scattering and
absorption of light of different wavelengths through the tissues of the patient 202 during
data ition. In one aspect, a correction term for light directionality may be measured
by ing data from one or more neous tissue phantoms and using a sensor
configuration in which no emission filters are present. The ratio of the signals detected by
the first light detector 222 (Bet!) and the signals detected by the second light detector 224
(Det2) measured are used to ine a coefficient Gex or Gem for signals obtained in
association with illumination by light at the excitation and emission wavelengths,
respectively. The coefficients are used to modify the signal ed by the first light
detector 222. In one aspect, the correction of the signals acquired in a homogeneous
medium by the first light detector 222 using the coefficients Gex or Gem render the signals
measured by the first and second detectors 222/224, as equivalent to within 20% of one
r. In other aspects, the correction of the signals acquired in a homogeneous medium
by the first light detector 222 using the ients Gex or Gem render the signals ed
by the first and second detectors 4 as equivalent to within about 10%, to within
about 5%, to within about 2%, and to within about 1%.
- detector non-linear response correction module
Referring again to A, the pre-processing subunit 1302 in this aspect
includes a detector non-linear response correction module 2210 configured to enable a
correction to variations in the detected signals associated with non-linear response of the
detectors. In this aspect, a calibration curve based on average data may be used to scale the
magnitude data obtained by the detectors 222/224.
-filter throughput temperature correction (emission) module
ing again to A, the pre-processing subunit 1302 in this aspect
includes a filter throughput temperature correction (emission) module 2212 configured to
enable a correction to variations in the ed signals ated with temperature-
dependent optical properties of the optical filter 244 associated with the second light
detector 224 during emission-wavelength illumination. In this aspect, the signals Det2
detected by the second light detector 224 may be corrected according to Eqn. (8):
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Det2 —Det2 (CemF,slopeT(T—Tnom))
Det2 = Eqn. (8)
CemF,nom
In various aspects, the signal Det2 ed by the second light detector
224 may be monitored while ambient temperature is cycled over a range including the
operating ature range or a large enough subset of the range to adequately determine
the ature-dependence of the emission filter. These data are acquired with the optical
filter 244 installed on the second light detector 224 from a homogeneous, non-fluorescent
phantom. Further, simultaneous measurements are monitored from the first light detector
222, and a ratio of the measurements Det2/Bet] is determined. The nominal filter
coefficient CemF,nom is calculated as the l ratio of Det2/Bet] obtained at a nominal
operating temperatureTnom. In this , the coefficient CemF,slopeT is obtained from the
slope of et] obtained over a range of ambient atures during emission-
wavelength illumination of the homogeneous, non-fluorescent phantom.
- tissue heterogeneity correction module
Referring again to A, the pre-processing subunit 1302 in this aspect
includes a tissue heterogeneity correction module 2216 configured to enable a correction to
variations in the detected signals associated with heterogeneity of the tissues intervening
between the first region 206 illuminated by light sources 218/220 and the second and third
regions 208/210 at which the light detectors 222/224 are positioned. In this aspect, the
signal Det] corrected for light directionality by the light ionality correction module
2208 and the signal Det2 corrected for filter effects by the filter throughput temperature
correction (emission) module 2212 are used to calculate Chetm, a coefficient to correct for
tissue heterogeneity, ing to Eqn. (9):
Chmm = Det2/Bet] Eqn. (9)
r throughput temperature correction (excitation) and signal decomposition module
ing again to A, the pre-processing subunit 1302 in this aspect
includes a filter throughput temperature correction (excitation) module and signal
decomposition module 2214 configured to enable a correction to ions in the detected
signals associated with temperature-dependent optical properties of the optical filter 244
associated with the second light detector 224 during excitation-wavelength illumination. In
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this aspect, e the emission filter is configured to block light at the excitation
wavelength, the filter throughput temperature correction (excitation) module and signal
osition module 2214 performs a correction to variance to the amount of excitation
light leakthrough due to temperature-related changes in the optical properties of the optical
filter 244. Further, the filter throughput temperature correction (excitation) module and
signal decomposition module 2214 enables corrections of the signals measured by the first
light detector 222 during excitation-wavelength illumination due to the presence of
fluorescence induced by the excitation-wavelength illumination superimposed over the
portion of the signal ated with the excitation-wavelength illumination.
In this aspect, the effects of temperature-dependent variation on rough
of excitation -wavelength by the optical filter 244 are calculated as expressed in Eqn. (10):
CexLT : CexLT,n0m + CexLT,slopeT(T _ Tnom) Eqn~ (10)
In this aspect, CexLT,n0m is calculated from the ratio of signals Bet] and
Del2 measured from a homogeneous, non-fluorescent phantom at the nominal operating
temperature Tnom during excitation-wavelength illumination. sl0peT is calculated as
the slope of the signal Del2 measured from a homogeneous, non-fluorescent phantom at a
range of operating temperatures T during emission-wavelength illumination.
In this aspect, the filter throughput temperature tion ation)
module and signal osition module 2214 further performs a signal extraction to
isolate ns of the detected signals associated with diffuse reflectance of the excitation-
wavelength illumination and fluorescence. DRexz which is the amount of excitation light
ent on the second light detector 224 in the absence of an l filter 244, is not
measurable, due to the presence of the optical filter 244. r, the signal Del] measured
by the first light detector 222 is a composite signal from both diffuse reflectance of the
excitation-wavelength illumination DRexl and fluorescence Flr] . CHetero is obtained using
the tissue heterogeneity correction module 2216 as described above. The underlying
signals are extracted by use of the following system of equations:
Detz : CexLTDRexZ + FlTZ Eqn. (11)
Detl : DRexl + FlT‘l Eqn. (12)
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Flrz : CHeteT-OFlrl Eqn. (13)
DRexZ : CHeteroDRexl Eqn~ (14)
In this aspect, Fir; is determined by solving the above system of equations
using only measurable s Dell and Det2 as demonstrated below:
Detz : CexLTCHeteroDRexl + F172 Eqn~ (15)
Detz : CexLTCHeter0(Det1 — Flrl) + Flrz Eqn. (16)
Detz : CexLTCHeteroDetl _ CexLTCHeteroFlrl + F172 Eqn~ (17)
Detz _ CexLTCHeteroDetl : FlT'2(1 _ CexLT) Eqn~ (18)
Detz—CexLTCHeteroDetl
F172 : Eqn (19)
1—CexLT
In this aspect, once Flr2 is obtained as described above, the other signals
Flr1,DRex1, and DRexz may be readily obtained through ion into the system of
equations (Eqns. (11) — (14)) presented above.
-fractionalphoton normalization module
Referring again to A, the pre-processing subunit 1302 in this aspect
includes a fractional photon normalization module 2218 configured to convert the detector
signals, after preprocessing as described above, into units of fractional photons for use in
subsequent background subtraction and intrinsic fluorescence correction algorithms as
described herein. In this aspect, the detector signals may be converted to urrent by
reversing the scaling associated with the ADC and the mpedance amplif1er used to
acquire the detected signals to obtain the s in units of photocurrents. Once
photocurrent is obtained, a detector responsivity supplied by the light detector’s
manufacturer is used to convert the detector urrents to units of Watts. The or
signals in Watts are then ratioed to the source power in Watts as measured by additional
light detectors 226 used to monitor the output of the light sources 218/220 to obtain the
number of fractional photons detected.
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- opticalpower correction module
Referring again to A and B, the pre-processing subunit
302a in this aspect es a fractional photon normalization module 2218/2218a
configured to convert the detector s, after preprocessing as bed above, into
units of fractional photons for use in subsequent background subtraction and sic
fluorescence correction algorithms as described herein. In this aspect, the detector signals
may be converted to photocurrent by reversing the scaling associated with the ADC and the
transimpedance amplifier used to acquire the ed signals to obtain the signals in units
of photocurrents. Once photocurrent is ed, a detector responsivity supplied by the
light or’s manufacturer is used to convert the detector photocurrents to units of Watts.
The detector signals in Watts are then ratioed to the source power in Watts as measured by
additional light detectors 226 used to monitor the output of the light sources 218/220 to
obtain the number of fractional photons detected.
- excitation light leakthrough subtraction module
Referring again to B, the pre-processing t 1302a in this aspect
includes a fractional photon normalization module 2222 configured to perform an
excitation leakthrough subtraction on the Flrmeas signal. To arrive at a fluorescence signal
due only to fluorescent photons (Flrphotons ), an excitation leakthrough subtraction is
performed. To remove the contribution of excitation light, the excitation leakthrough is
taken to be a fraction of the diffuse reflectance excitation (DRexmeas) signal, where a
universal calibration factor, CExLT, determines the fraction of the signal to subtract from
Flrmeasas expressed below:
EXLT : CEXLT * DR
exmeas
where CExLTis a calibration factor that is obtained by computing the ratio between the
tion light detected by both detectors on a non-fluorescing optical m as
described below:
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Flrmeas
CExLT — DR—_
exmeas
This signal is then subtracted from sto provide a fluorescence signal
due only to fluorescent photons as expressed below:
tons : Flrmeas _ EXLT
-fluorescence light leakthrough subtraction module
Referring again to B, the pre-processing subunit 1302a in this aspect
includes a fluorescence light leakthrough subtraction module 2224a configured to perform
a fluorescence leakthrough subtraction on the Flrmeas signal. To obtain the diffuse
reflectance, defined herein as the excitation signal due to only excitation photons
(DRexphotonS ), a fluorescence leakthrough subtraction is performed. To remove the
fluorescence leakthrough, a calibration factor, CFlrLT: was determined based on the
relationship between the amount of fluorescence leakthrough observed on a se of
human subject data and tissue heterogeneity as measured by the onship between the
diffuse reflectance, emission signals (Dilifilt). The relationship is a linear relation as
expressed below:
DRem
C 1 < : * —
“T” p p2
lt>+
where p1 and p2 are approximately 0.61 and 0.01, respectively, in one aspect, as
determined by the mentioned relationship. In r aspect, p1 and p2 may assume
any other value without limitation as defined by the above relationship.
The DRexphotonS signal is then calculated by cting this fraction of
measured fluorescence from the e reflectance excitation signal, as follows:
DRexphotons : DRexmeaS _ Flrmeas * CFlrLT
b) baseline subtraction subunit
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Referring again to , the processing unit 236 further includes a
ne subtraction subunit 1304. In an aspect, the ne subtraction subunit 1304
cts a baseline signal from the light detector measurements to correct for the effects of
orescence and light leakage. The baseline period, as used herein, refers to an initial
time period of measurements obtained prior to injection of the exogenous fluorescent
agent. During the baseline period, the fluorescence signal measured by the system 200 may
be assumed to be associated with tissue autofluorescence and/or excitation light from the
LED light sources 218/220 leaking h the optical filter 244 of the second light
detector 224. In an aspect, the average signal measured during the baseline period, referred
to herein as a ne signal, may be subtracted from subsequent cence
measurements to yield a measurement associated solely with the fluorescence produced by
the exogenous fluorescent agent within the tissues of the patient.
In another aspect, the corrections for excitation light leakthrough and
autofluorescence may be implemented in cooperation with the background correction
subunit 1306. In this other aspect, rather than subtracting an average signal ed
during the baseline period, the background tion subunit 1306 may dynamically
calculate the effects of tion light leak-through and orescence at each data
acquisition cycle. As a result, subtraction of the effects of excitation light leak-through
may be performed prior to the diffuse reflectance correction described herein below, and a
subtraction of the effects of autofluorescence may be updated at each data acquisition cycle
by the background correction subunit 1306.
c) background correction subunit
In an aspect, the background correction t 1306 may correct the
measured fluorescence data to remove the effects of changes to the optical ties
(absorption and scattering) of the tissues of the patient 202 during monitoring of renal
extraction of an exogenous fluorescent agent within the tissues of a patient. As described
herein above, the optical properties of the tissues may change due to any one or more
factors including, but not limited to: vasodilation, vasoconstriction, oxygen saturation,
hydration, edema, and any other suitable factor within the region of interest monitored by
the system, associated with changes in the concentrations of endogenous fluorophores such
as hemoglobin, en, and melanin.
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In one aspect, the background correction subunit 1306 may determine the
intrinsic autofluorescence (IFauw) signal, representing the emission-wavelength light
emitted by endogenous fluorophores within the tissues of the patient during data
acquisition. In this aspect, the [Fam signal is obtained from the mean or median of [Fbkmd-
(the background intrinsic fluorescence data prior to agent ion). The IFbkmd signal is
found as s:
Flrbkrnd
IFbkrnd =
DRex DRembkx bkmDRemFiltbkmFilt
where the coefficients bkx, bkm, and kmFilt are found via a global error e method.
In one aspect, the values of the powers used in the equation above are
determined empirically using a global error surface method. The method in this aspect
includes selecting ranges of values for each of the powers (bkx, bkm, bkmFill) for each of
the diffuse reflectance signals (DRex, DRem, ‘ltgrgd) selected by a user. In various
aspects, the ranges of values for each of the powers may be influenced by any one or more
of a variety of factors ing, but not limited to: the design of the system 200, including
the design of the sensor head 204, the properties of the selected exogenous fluorescent
agent such as excitatory/emission wavelengths, absorption ncy, emission efficiency,
and concentration of initial dose in the patient’s tissues, the species of the patient 202 and
corresponding concentrations of endogenous chromophores, the position of the sensor head
204 on the patient 202, and any other relevant factor.
In one aspect, the method may include choosing a wide range for each
coefficient (bkx, bkm, l) and t a broad search. The error surfaces from this
broad search may be analyzed to locate wells in the error surface and the associated ranges
for each of the coefficients. The method in this one aspect includes adapting the ranges of
each coefficient to include the regions from the broad search within which wells in the
error surface were observed and repeating the analysis. This method may be iterated until a
suitably fine resolution is achieved that is capable of tely capturing the minimum
error.
Step sizes may be selected at 1404 for the ranges of values selected for each
power (bkx, bkm, bkmFill). In an aspect, the step size for each factor may be selected based
on any one or more of at least several factors including, but not limited to: the anticipated
W0 2018/140984 PCT/U82018/016053
sensitivity of the [F values calculated above to s in each factor; a suitable total
number of combinations of powers used to calculate IF considered factors including
available computational resources, acceptable data processing times, or any other relevant
s; and any other suitable criterion for step size.
In various aspects, the step sizes may be the same value for all powers (bkx,
bkm, bkmFill). By way of non-limiting example, the step size for all powers may be 0.5. In
various other aspects, the step sizes may be constant for all values of a single power (bkx,
bkm, bkmFill), but the step sizes selected for each power may be different between
different powers. By way of non-limiting example, the selected step size for bkx may be
0.01 and the selected step size for bkm and bkmFl'lZ may be 0.6. In s additional
aspects, the step size within one or more of the powers may vary within the range of values
for each power. In these various additional aspects, the step size may be reduced within
subranges of values for a power for which the IF calculated above is predicted to be more
sensitive to small changes in that power. Non-limiting examples of suitable varying step
sizes within a range of values for a single power include: different step sizes selected by a
user, random step sizes, a linear increase and/or se in step size, a non-linear
distribution of different step sizes such as a logarithmic distribution, an exponential
distribution, or any other suitable non-linear distribution of step sizes.
The selected ranges of exponents, together with the ed step sizes, may
be used to form vectors of potential values of bkx, bkm, bkmFill. For each combination of
exponents amongst all vectors, [F is calculated from the ements Flr, DRex, DRem,
and DRng’l‘ltgrgd using the above equation. For each combination of exponents, a plurality of
[F values are calculated in which each IF value corresponds to one of the data acquisition
cycles By way of non-limiting example, using the vectors of potential exponents listed
herein above, a total of 405 (5*9*9) pluralities ofIF s would be calculated.
In an aspect, the plurality of combinations of potential exponents may be
evaluated to select one combination of nts from the plurality to assign for use in
subsequent diffuse reflectance corrections calculated using the above on. An te
of error of the corrected Flr signal data (i.e. IF signal data ated using the above
equation may be calculated. Any estimate of error may be calculated including, but not
limited to, a quantity d to residuals of the IF signal data relative to a curve fit of the IF
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signal data. Any type of known curve-fitting method may be used to curve-fit the IF signal
data ing, but not limited to, a single-exponential curve fit. Without being limited to
any particular theory, it is thought that the rate of clearance of an exogenous fluorescent
agent, such as lVfl3-102, from the kidneys is expected to be a constant exponential decay
characterized by the renal decay time constant RDTC.
Intrinsic autofluorescence (IFauw) is then simply the mean or median of
IFbkmd~
The autofluorescence signal, Flrauw, is then ted by performing the inverse
background diffuse reflectance correction, as s:
Flam = IFauw * (DRsrm/ DREEx/DRé’fll-fi”
This autofluorescence signal, Flrauw is then removed from the measured fluorescence
signal, Flr, to determine the agent intrinsic fluorescence (IFagem) cally
representing the emission-wavelength light d by the exogenous fluorescent agent.
Without being limited to any particular theory, the cence
measurements obtained by the system 200 that are used to determine renal function include
emission-wavelength photons that are detected by the second (filtered) light detector 224.
These emission-wavelength photons are emitted by the exogenous fluorescence agent
introduced into the tissues of the patient in response to illumination by excitation-
wavelength photons. The emission-wavelength photons travel from the cence source
(i.e. the exogenous fluorescence agent) to the second (filtered) light detector 224 through
third region 210 of the patient’s skin. However, the emission-wavelength light that is
detected by the second (filtered) light detector 224 may also include orescence
emitted by nous chromophores such as keratin and collagen within the tissues of the
patient, as well as hrough of excitatory-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 produced by the first light source 218
and are directed into the first region 206 of the patient’s skin. If the optical properties of the
patient’s skin (scattering and/or absorption) varies over the time interval at which the
or data used to determine renal function is acquired (i.e. from a few hours to about 24
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hours or more), the accuracy of the fluorescence measurements may be impacted, as
discussed previously above.
During each measurement cycle in an aspect, the system 200 may direct
light into the first region 206 of the patient’s skin with a pulse of emission-wavelength light
and a pulse of tion-wavelength light in an alternating series and may detect all light
emerging from the second region of the patients skin using the first (unfiltered) light
detector 222 and a portion of the light emerging from the third region 210 of the patient’s
skin using the second (filtered) light detector 224. The light intensity detected by each
ation of tion and emission ngth illumination of the first region 206 and
detection by the unfiltered/filtered light detectors 222/224 contain information not only
about the tration of the exogenous fluorescent agent in the t’s tissues, but also
information about the optical properties of the patient’s skin.
Table 2.‘ Light Detector Measurements After Temperature and Power Fluctuation
Corrections
Illumination First (Reference) Second (Primary)
wavelength Light or Light Detector
Unfiltered Filtered
The primary measurement of fluorescence is Flrmeas the
, intensity of
fluorescent light measured at the filtered detector.
The diffuse reflectance measurement Flrmeas represents the propagation of
photons to the non-filtered arm and is composed primarily of tion photons.
DRemand DRemlfl-lteredrepresent the propagation of emission-only photons.
Referring to Table 2, light intensity measured by the second ed) light
detector 224 during illumination by the excitation-wavelength light captures the raw
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intensity of light emitted by the exogenous fluorescent agents (Flrmeas) prior to any
corrections for tissue optical properties in various aspects. After baseline subtraction
corrections as described herein previously, the emission-wavelength light contained in
Flrmeasis assumed to originate predominantly from the exogenous fluorescent agent, with
only minor butions due to auto-fluorescence by endogenous fluorophores, and is
ore termed Flragent. In an aspect, if no change in the optical properties of the
patient’s skin is assumed, all autofluorescence butions would be subtracted off during
the baseline correction described herein above.
However, if the optical properties of the t’s skin change during the
acquisition of data, slightly more or less of the orescence may emerge from the
patient’s skin at the emission wavelength, thereby introducing uncertainty into the accuracy
of the background ction correction performed previously. In addition, varied skin
optical properties may further alter the ity of light at the excitation wavelength
reaching the exogenous fluorescent agent, thereby altering the amount of energy absorbed
by the exogenous fluorescent agent and the intensity of induced fluorescence from the
exogenous fluorescent emitted in response to illumination by the excitation-wavelength
light. In various aspects, the remaining three light measurements enable monitoring of the
optical ties of the patient’s skin and provide data that may be used to adjust for any
changes in the optical properties of the patient’s skin including the effects of
autofluorescence and excitatory-wavelength light bleed-through.
Referring again to Table 2, light intensity measured by the first (unfiltered
reference) light detector 222 during illumination by excitation-wavelength light captures a
measure of the diffuse ance of excitation-wavelength light propagated through the
t’s skin (DRex). Although the first light detector 222 is configured to detect both
excitation-wavelength and emission-wavelength light, the intensity of the excitation-
wavelength light is orders of magnitude higher than the ity of the emission-
wavelength light as a result of the lower efficiency of producing light via fluorescence. In
various aspects, the proportion of emission-wavelength light within DRexis assumed to be
negligible. In other aspects, the proportion of emission-wavelength light within DRex is
estimated and subtracted. t being limited to any particular theory, because the
ity of the excitation-wavelength light directed into the patient’s skin is d to be
relatively constant with ible losses due to absorption by the exogenous fluorescent
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agent, and is subject to power corrections as described herein usly, DRex serves as a
benchmark measurement to assess changes in the optical properties of the patient’s skin
with respect to the excitation-wavelength light.
Light intensity measured by the first ered reference) light detector 222
during illumination by emission-wavelength light captures a measure of the diffuse
ance of emission-wavelength light propagated through the patient’s skin (DRem).
Without being limited to any particular theory, because the exogenous fluorescent agent is
not d to emit emission-wavelength light due to the absence of excitation-wavelength
illumination during this phase of the data acquisition cycle, and because the intensity of the
emission-wavelength light directed into the patient’s skin is relatively constant and subject
to power corrections as described herein previously, DRex serves as a benchmark
measurement to assess changes in the optical properties of the patient’s skin with respect to
the emission-wavelength light.
Light intensity measured by the second ed) light or 224 during
illumination by emission-wavelength light captures a second measure of the diffuse
reflectance of emission-wavelength light propagated through the patient’s skin
(DRgmflltgrgd). In one aspect, DRng’l‘ltgrgd is subject to the same assumptions as DRem as
described herein above. In addition, tgrgd provides a means of assessing
heterogeneity of the tissue’s optical properties. Because DRemflmgd is measured by the
second light detector 224 configured to detect light emerging from the patient’s skin at the
third region 210 (see , the intensity of light measured in DRgmflltgrgd has propagated
along an l path through the skin of the patient that is different from the optical path
led by the light measured in DRem. Without being limited to any particular theory,
e the distances of the first detector aperture 1004 and second light aperture 2006
h which light is delivered to the first and second light detectors 222/224, respectively
are designed to be equidistant from the light delivery aperture 1002 (see ), any
differences between DRemflltgrgd and DRem are assumed to be a result of heterogeneity on the
optical properties of the skin traversed by the two ent optical paths.
excitation-wavelength light leak-through correction
In one aspect, DRexmeas serves as a basis for the estimation of leak-through
of excitatory-wavelength light into the second (filtered reference) light detector 224 used as
W0 40984 PCT/U82018/016053
part of the method of ng the effects of variation in background signal described
herein. Without being limited to any particular theory, it is assumed that the amount of
leak-through of excitatory-wavelength light into the second (filtered reference) light
detector 224 is tional to the DRM signal, and that this proportion is influenced
exclusively by device-related factors, rather than factors related to the optical properties of
the patient’s skin. As a result, the proportion of the DRM signal representing leak-through
light is assumed to be constant, as described herein below.
In one aspect, the excitation-wavelength light leak-through (ExLI) included
within the raw fluorescence signal (FIt") is assumed to be a constant fraction CExLT of the
DRexmeas signal according to Eqn. (21):
EXLT : CEXLT * DR
exmeas Eqn. (21)
where CEx” is a sensor-head specific calibration factor.
In one aspect, CEx” is obtained by computing the ratio between the
excitation light detected by first and second light detectors 4 (Dell/Del2) on a non-
fiuorescing optical phantom ing to Eqn. (22):
Det2 Flrmeas
CExLT = _
_ Eqn. (22)
Detl DRexmeas
In another aspect, the excitation light reaching the filtered detector is
d to be different than the light reaching the non-filtered detector due to tissue
heterogeneity. In this aspect, the ratio of the emission-wavelength light at each detector is
used to t for this heterogeneity.
In various aspects, the calibration factor CExLT may be specific to an
individual sensor head 204 or CEx” may be applicable to all sensor heads 204 of a system
200 depending on various factors including, but not limited to, cturing nces.
In an aspect, if the system 200 is used to obtainCExLT, s and DRexmm are from a
non-fluorescent, homogeneous phantom in the context of the system 200 as described
herein above. It is to be noted that Eqn. (22) assumes that the tissue monitored by the
system 200 is homogeneous.
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In an aspect, the excitation-wavelength light leak-through (ExLT)
determined by Eqn. (21) may be subtracted from the raw fluorescence signal (Flrmeas) to
obtain a corrected fluorescence signal ns as described in Eqn. (23):
Flrphowns = Flrmeas — ExLT Eqn. (23)
A is a graph of a raw fluorescence signal (Flrmeas, blue line) and the
corresponding excitation-wavelength light hrough (ExLT, red line) determined using
Eqn. (23) obtained by a system 200 in one aspect before and after the injection of an
exogenous fluorescence agent. As illustrated in A, the ExLT signal varies over the
course of data ition. B is a graph comparing the raw fluorescence signal
(Flrmeas, blue line) and the fluorescence signal with the excitation-wavelength light leakthrough
removed (FlrphotonST, green line) as described herein above in Eqn. (23).
In one aspect, the raw fluorescence signals s are first corrected to
remove the effects of tion-wavelength light leak-through using Eqn. (23). In this
aspect, subsequent corrections to remove the effects of autofluorescence are implemented
using the corrected fluorescence signal Flrphotonsas a basis as described herein below.
fluorescence leak-through correction
Without being limited to any particular theory, the light detected by the
unflltered light detector during illumination by light at the excitatory wavelength is a
mixture of diffuse reflectance of the excitation wavelength light and light from agent
fluorescence. In one aspect, the diffuse reflectance is d to be sufficiently more
intense than the fluorescence such that the contribution of fluorescence to the unflltered
detector ement was negligible.
In another aspect, the contribution of fluorescence to the red detector
measurement may be non-negligible. By way of non-limiting e, is a graph
showing DRexmeas and Flrmeas over a full day in the absence of administration of an
ous fluorescent agent. However, as illustrated in , the DRexmm signal
occasionally showed leak h of fluorescence, as evidenced by a correlated signal rise
after agent administration into the patient’ s bloodstream.
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In one aspect, the portion of the e ance excitation signals due
only to excitation photons are removed from the raw DRexmeassignal according to Eqn.
(24):
DRexphotons : DRexmeaS _ Flrmeas * CFlrLT Eqn~ (24)
In various aspects, the coefficient CFM” is empirically determined using the
relationship between a measured amount of fluorescence leak-through (Flrleakthmugh) on
DRemFilt
the DRexmeaS signal, in relation to tissue heterogeneity as expressed by the ratio
DRem
(see discussion below). In one aspect, the measurements may be obtained from a plurality
of subjects. By way of miting example, is a graph summarizing a
DRem
relationship between empirically determined Flrleakthmugh and d from a
DRemFth
database of 33 ts. In this aspect, this empirically-derived relationship was checked
on multiple patient datasets and found to be consistent. The correction coefficient CFlrLT
was set to incorporate the relationship between tissue heterogeneity and amount of
fluorescence leak through, as defined below:
CFlrLT = P1 * (—DRW::.”) + p2 Eqn. (25)
In one , Eqn. (25) includes p1 =0.6l38 and p2 = 0.01095, as
determined by a bisquares weighted linear fit to the relationship illustrated in .
In another aspect, CFITLT is determined by obtaining measurements on
optical phantoms provided with increasing fluorescence concentrations, where the only
change signals is due to tration of exogenous fluorescence agent concentration.
ion of fluorescence and excitation wavelength diffuse reflectance
In various aspects, the number of photons due to either DRex or Flr on either
the filtered or unfiltered detector depends on light directionality and the gain of each
detector at the detected wavelength, as shown below:
DRexmeaS = A1 * DRexphotonS + B1 * Flrphotons Eqn. (26)
s : A2 * otons + 32 * Flrphotons Eqn~ (27)
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where the coefficients A1, A 2, B1, and B2 include a directionality and gain factor. By way
of non-limiting example, A1 may be provided in the form of Eqn. (28):
A1 = d4505PM1 * GSPM1@450 Eqn. (28)
where_d4505PM1 and GSPM1@450are directionality and gain factors of a detector SPMI at
an illumination wavelength of 450 nm.
In one aspect, the photon s may be isolated as expressed in Eqns. (29)
and (30):
B A A
Bl (B—i _ (Ti) tons : Flrmeas _ 1T: Rexmeas Eqn. (29)
A B B
A2 (I: — 57:) own, = DRexmm — B—:Flrmeas Eqn. (30)
In various aspects, the constant terms in front of the photon signals, such as
B A
B1 (B—Z—A—Z) . . .
are not , because the renal functlon. rs as d1sclosed here1n
1 1
measure rates of change of intrinsic fluorescence (IF) as expressed by Eqn. (31):
IF = C0 + Cle't/T —> log(IF) = log(C1) —% Eqn. (31)
In one aspect, the terms % (or CM”) and g (or CFlrLT) are determined
1 2
experimentally to isolate Flrphotons and DRexphotonS, respectively, as described above.
autofluorescence correction
In various aspects, the method of correcting the measured fluorescence to
remove the time-varying effects of background may r include removing the effects of
autofluorescence in addition to removing the effects of excitation-wavelength leak-though.
Autofluorescence, as used herein, refers to the emission-wavelength light produced by
endogenous chromophores, such as keratin and collagen, in response to illumination by
excitation-wavelength light. In s aspects, autofluorescence may vary over the course
of acquiring fluorescence measurements using the systems and s described herein.
Without being limited to any particular theory, changes in the optical properties of the
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patient’s skin, such as s in the concentration of chromophores such as hemoglobin
and/or melanin may cause variation in autofluorescence levels.
is a graph summarizing the measurements of raw fluorescence
(Flrmeas, blue line) obtained during the background interval, defined herein as the interval
prior to injection of the exogenous cent agent into the patient 202, when the
patient’s tissues are assumed to contain no exogenous cent agents. Also shown in
is the signal resulting from the removal of the effects of hrough of excitatory-
wavelength light (ExLI) from Flrmeas using Eqn. (5) as described herein above. The
remaining fluorescence signal detected during the background interval, shown as a green
line in , may be assumed to be attributable to autofluorescence in various aspects.
In one aspect, the intrinsic autofluorescence (IFauw), defined here as the
measured fluorescence at the emission wavelength attributable only to emission by
endogenous chromophores, such as n and collagen, may be calculated as the median
value of the corrected fluorescence signal Flrphotons (see Eqn. 23) obtained during the
background interval according to Eqn. (32).
IFAutO = median(Flrph0t0nS(1: endBackgroundD Eqn. (32)
where endBackground is the index of the data acquisition in the dataset ponding to
the end of the background interval just prior to injection of the exogenous fluorescent
agent.
In an aspect, the orescence may be assumed to be relatively stable
throughout the entire data acquisition process, including the interval ing injection of
an exogenous fluorescent agent. In this aspect, the effect of autofluorescence may be
removed by subtracting the IFAutO value obtained in Eqn. (32) from the corrected
fluorescence signal Flrphotonsas expressed in Eqn. (33):
[Fagent : Flrphotons _ IFAuto Eqn~ (33)
where [Fagem denotes the intrinsic cence specifically representing the emission-
wavelength light emitted by the exogenous fluorescent agent.
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A is a graph summarizing various measurements obtained during the
background al: raw fluorescence (Flrmeas), DRexmeas (red line), DRem (orange line),
and DRemfltmd (purple line). In on, the intrinsic autofluorescence (IFauw, green line)
calculated using Eqn. (32) is also shown in A. During the background interval
shown in A, all quantities were vely stable in value.
B is a graph summarizing the diffuse ion ements
shown in A: DRexmeas (red line), DRem (orange line), and DRgmflltgrgd (purple line).
Over the course of obtaining fluorescence measurements after injection of the exogenous
fluorescent agent (i.e. after a time of about 9:07 as shown in ), the e reflection
measurements decrease signif1cantly, indicating that the optical properties of the patient’s
skin, which impact the measured signal from autofluorescence, may also change during this
time period.
In an additional aspect, diffuse reflectance measurements may be used to
project the underlying autofluorescence signal for the full ement period, thereby
accounting for changes in the l properties of the patient’s skin over the full course of
data measurements. In one , diffuse reflection measurements may be used to scale
the corrected fluorescence signal Flrphotons to account for changes in the optical properties
of the patient’s skin, resulting in an intrinsic fluorescence. In this aspect, to t the
fluorescence measurements obtained after the injection of the ous fluorescence
agent, the intrinsic autofluorescence (IFauw) calculated from Eqn. (32) may be subtracted
from the combined sic fluorescence [FAggnmndAutg obtained from Eqn. (33), as
expressed in Eqn. (34):
[Fagent : IFAgentAndAuto _ IFAuto Eqn~ (34)
In one aspect, the background correction subunit 1306 may enable a
background correction method 2000 as summarized in the block diagram of . The
method 2000 may include performing a correction at 2002 to remove the s of the
leak-through of emission-wavelength light into the second (filtered reference) light detector
224 as described in Eqns. (29), (30), and (31) above. The method 2000 may further include
estimating the level of autofluorescence (IFauw) at 2004 from an analysis of the
measurements obtained during the background interval as described in Eqn. (32) above.
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The method 2000 may additionally include performing a correction at 2006 to remove the
effects of autofluorescence from the fluorescence measurements as described in Eqn. (33)
above. In effect, the autofluorescence signal [Fame is projected forward into subsequent
fluorescence measurements and is removed at 2006. The intrinsic fluorescence [Fagem
resulting from the removal of background effects from the raw fluorescence measurements
may transformed by the RDTC calculation subunit 1310 as described herein below into a
parameter including, but not d to, glomerular ion rate (GFR) and/or renal decay
time constant (RDTC) representing renal function.
e) fault detection subunit
Referring again to , the sing unit 236 of the controller 212
may further include a fault detection subunit 1312 configured to monitor the function of the
light sources 218/220 and light detectors 222/224 and to inform the user of any
irregularities of any detected faults within the system 200 via the y unit 216. In
various s, the fault detection subunit 1312 may enable the basic identification of fault
and notice states by examining the signal levels ed from the light sources 218/220
and light detectors 4 and associated additional temperature sensors 228 and
additional light detectors 226 of the sensor head 204 (see . In various aspects, the
signal magnitudes (see Eqn. (1)) and average signals may be used to determine the peak
and nadir levels of the modulation of the LED light sources 218/220. The nadir of the
signal, defined herein as the average signal minus half the peak-to-peak signal, may be
used to monitor ambient light levels in one . t being limited to any particular
theory, additional contributions to the nadir levels of the modulated signals, such as
er DC offset, may be neglected as small and constant relative to the contributions of
ambient light leakage. In an aspect, if the detected ambient light levels register in excess of
about one quarter of the high-speed ADC 1102 range at low or amplifler gain, an
ambient light notice is issued to the user via the display unit 216.
In various other aspects, saturation of the light detectors 4 detectors
may also be monitored by the fault detection subunit 1312. In these other aspects, the
saturation may be monitored by calculating the peak value of the , defined herein as
the average signal value plus half the peak-to-peak signal. If the signal’s peak value falls
within is within 5% of saturation of the ADC range, the fault detection subunit 1312 may
W0 40984 PCT/U82018/016053
issue a saturation notice to the user via the display unit 216. If saturation event is detected
by the fault detection subunit 1312, the ambient light level may then be checked to
determine if the saturation event is associated with ambient light saturation, defined herein
as a saturation event occurring concurrently with an ambient light notice as described
herein above. If an ambient light saturation event is detected, the fault detection subunit
1312 issues an ambient light saturation notice to the user via the display unit 216, and data
ition by the acquisition unit 234 is continued in this notice state to allow the user to
resolve the ion. If a saturation event is detected that is not associated with an excess
of ambient light, the fault detection subunit 1312 may signal the light detector control unit
232 to perform an adjustment of detector gain and/or may signal the light source control
unit 230 to perform an adjustment to the LED t source 1126 to adjust LED intensity.
In various aspects, the fault detection subunit 1312 issues a notification to the user via the
display unit to report either the ambient light saturation event, or the saturation event not
associated with an excess of ambient light. In some aspects, if a tion event is
detected, but the automatic gain adjustment has been disabled by a user when the system
200 is configured in the ering Mode as described herein above, the user is also
notified via the display unit.
e) post-agent administration selection subunit
Referring again to , the processing unit 236 may further include a
post-agent administration subunit 1308 configured to automatically identify the portion of
the measurement data set that corresponds to an gent administration region, as
described herein below.
is a graph of fluorescence ements obtained from a t
over a period of about 10 hours after injection of an exogenous cence agent such as
1Vfl3-102 after a pre-injection period 2102 of about 3 hours. Referring to , the pre-
injection/baseline period 2102 is characterized by a relatively low and stable fiuorescence
level, likely due the absence of exogenous fiuorescent agent in the blood of the patient.
After the injection 2103 of the exogenous fiuorescence agent, the fluorescence
measurements exhibit a sharp increase 2106 to a peak concentration 2108, followed by a
relatively smooth exponential decrease back to ound fiuorescence levels at the
s eliminate the exogenous fiuorescence agent from the blood of the patient.
W0 2018/140984 PCT/U82018/016053
Without being limited to any particular theory, it is thought the ed exogenous
fluorescence agent is likely well-mixed after an amount of time in the exponential
concentration decrease has elapsed.
Referring again to , after an exogenous fluorescent agent, such as
1Vfl3-102, is injected into the bloodstream of a patient, the exogenous fluorescent agent
undergoes an equilibration period of diffusion from the bloodstream into the rest of the
extracellular tissues of the patient. After agent ion 2103, the temporal profile of the
fluorescence signal IF may be characterized as a two-exponential signal profile described
by Eqn. (35):
IFpre—equilibration : C0 + 616—”?1 + Cze—t/TZ Eqn~ (35)
in which C0 is the baseline signal that is typically removed by baseline subtraction as
described herein above.
Referring again to , once the diffusion of the exogenous fluorescent
agent into the extracellular tissues of the patient reaches a quasi-steady state condition,
quilibration 2110 is achieved and the fluorescence signal may be characterized as a
linear decay. Without being limited to any ular theory, the quilibration region
2110 of the measurement data set is assumed to be terized as a temporal region of
the IF dataset that when log-transformed, is well-described by a linear equation. In one
aspect, the post-equilibration region is escribed bed by Eqn. (3 6):
IFpost—equilibration : C0 + Cle—t/T Eqn~ (36)
In an aspect, the post-agent administration selection subunit 1308 may
identify the quilibration period 2110 automatically by performing a single-exponent
curve fit at different portions of the IF data set and analyzing the associated curve fitting
errors for each of the different portions. In various aspects, the gent administration
selection subunit 1308 may select the earliest-occurring portion of the IF data set in which
the curve-flt error associated with a single-exponent curve fit falls below a threshold value
as the initial post-equilibration portion of the IF data set suitable for data correction and
analysis as described herein above. Any is method le for comparing curve-fit
errors association with single-exponential curve fits of different portions of the IF data set
W0 2018/140984
may be used in the post-agent administration selection subunit 1308 including, but not
limited to, linear curve-fitting portions of the IF data set falling within overlapping or non-
pping data windows and comparing the curve-fit errors of the ponding data
windows. In an aspect, the post-agent administration selection subunit 1308 may produce
at least one signal configured to signal the time range within the IF data set corresponding
to the post-equilibration period 2110 to the RDTC calculation subunit 1310 to enable the
selection of a suitable portion of the IF data set to correct and analyze as disclosed herein.
In r aspect, a linear fit and a 2-exponential fit to the IF data may be
compared. In this other aspect, equilibration may be identified as complete once the fitting
error is equivalent (corrected for the extra degrees of freedom in the nential fit).
fl RDTC calculation subunit
In various aspects, the system 200 is configured to transform the various
measurements from the light detectors 222/224 and associated light sources 0 and
other l and light sensors into a corrected intrinsic cent (IF) signal
corresponding to the detected fluorescence attributable solely to emission of fluorescence
by the exogenous fluorescent agent at the emission wavelength in response to illumination
by light at the tory wavelength. In various aspects, the exponential decrease of the IF
signals during the post- agent administration portion of the IF data set may be analyzed to
monitor and quantify renal function.
In one aspect, the exponential decrease of the IF signals during the post-
agent stration portion of the IF data set may be transformed into a glomerular
filtration rate (GFR) configured to quantify renal function. In another aspect, the
exponential decrease of the IF signals during the post-equilibration portion of the IF data
set may be ormed into a renal decay time nt , also configured to
quantify renal function. In another aspect, the exponential decrease of the IF signals during
the post-equilibration portion of the IF data set may be transformed into a renal decay rate,
also configured to quantify renal function.
Referring again to , the processing unit 236 may r include an
RDTC calculation subunit 1310 configured to automatically orm the IF signals into a
renal decay time constant (RDTC). As used herein, renal decay time constant (RDTC) is
W0 2018/140984 PCT/U82018/016053
defined as the time constant associated with the post-equilibration single-exponential decay
described in Eqn. (36) herein above. In one aspect, after accurate baseline subtraction by
the baseline subtraction subunit 1304, the renal decay time constant I may be calculated by
performing a linear regression on the log-transformed IF signal data (log (IF)), as described
in Eqn. (37):
log(IF) = ) — it Eqn. (37)
In various aspects the RDTC ation subunit 1310 may produce signals
configured to produce a display of the calculated RDTC using the display unit 216. The
display of the calculated RDTC may be provided to the display unit 216 in any suitable
format ing, but not limited to: a graph of RDTC as a function of time, a single
discrete RDTC value, a table of RDTC values as a function of time, a coded y
or other graphical representation configured to specific whether the calculated RDTC may
be classified as normal/healthy, abnormal, high, low, and any other suitable classification.
In various other aspects, any of the graphical formats described above may be continuously
or non-continuously updates as additional data is obtained and analyzed. In one aspect, the
RDTC calculation subunit 1310 may calculate RDTC as described herein above within
non-overlapping and/or overlapping windows within the IF data set.
In another aspect, the RDTC calculation t 1310 may t RDTC
into ular ion rate (GFR) using known methods. In this aspect, RDTC may be
inverted and multiplied by a slope, resulting in cGFR, a prediction of GFR that may be
corrected for body size (e.g. body surface area, or volume of distribution).
v) Memory
Referring again to the controller 212 of the system 200 may further
include a memory 242 red to facilitate data storage in the system 200. In some
ments, the memory 242 es a plurality of storage components such as, but not
limited to, a hard disk drive, flash memory, random access memory, and a magnetic or
optical disk. Alternatively or additionally, the memory 242 may include remote storage
such 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, cause the at least
one processor to perform any of the functions of the controller 212 described above. In one
W0 2018/140984 PCT/U82018/016053
implementation, the memory 242 may be or n 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 e area network or other configurations. A computer program product can
be tangibly embodied in an information carrier. The computer program product may also
contain instructions that, when executed, perform one or more functions, such as those
described herein. The information carrier may be a ansitory computer- or e-
readable , such as the memory 242 or memory on the processor 238.
In various s, the system 200 may record raw measurements and
processed data to a series of files. Each file may contain a header, which contains
information about the operator, instrument, and session. Each experimental session records
a set of files into a separate folder for each sensor head used in that session. The raw data
file may contains in-phase, quadrature, and average measurements from the detectors and
monitors during the active periods of both the excitation wavelength and the on
wavelength LEDs, along with the gain settings of the LEDs and detectors at the time of
data acquisition.
In various other aspects, the processed data file may contain the
fluorescence and diffuse reflectance measurements after magnitude calculation and
correction for the monitor readings, along with the gain settings of the LEDs and detectors.
The intrinsic fluorescence data file may contain the intrinsic fluorescence measurements
resulting from the diffuse reflectance correction of the raw fluorescence signals. The GFR
file may n the calculated GFR as a function of time, classified to te whether
post-equilibration has occurred, along with confidence bounds. The telemetry file may
n the temperature and e measurements. The event record file may contain both
user and automatically generated event records.
v1) GUI Unit
ing again to the controller 212 may include a GUI unit 240
configured to e a plurality of signals ng various measured and transformed
data from other units of the system in various aspects. In addition, the GUI unit may be
configured to produce signals configured to operate the display unit 216 in order to display
W0 2018/140984
data, frames, forms, and/or any other communications of information between the user and
the system 200.
viz) Processor
Referring again to the controller 212 may further include a processor
238. The processor 238 may include any type of conventional processor, microprocessor,
or processing logic that interprets and executes instructions. The processor 238 may be
configured to process instructions for execution within the controller 212, including
instructions stored in the memory 242 to display graphical information for a GUI on an
al input/output device, such as display unit 216 coupled to a high speed interface. In
other implementations, multiple processors and/or multiple buses may be used, as
appropriate, along with multiple memories and types of memory. Also, multiple controllers
212 may be ted, with each device providing portions of the necessary operations to
enable the functions of the system 200. In some ments, the processor 238 may
e the acquisition unit 234, the light detector control unit 232, the light source control
unit 230, and/or the processing unit 236.
As used herein, a processor such as the processor 238 may include any
programmable system including s using micro-controllers, d instruction set
ts (RISC), application specific integrated circuits (ASICs), logic circuits, and any
other circuit or processor capable of executing the functions described herein. The above
examples are example only, and are thus not intended to limit in any way the definition
and/or meaning of the term “processor.”
As described herein, computing devices and computer systems include a
processor and a memory. However, any processor in a er device ed to herein
may also refer to one or more processors wherein the processor may be in one computing
device or a plurality of computing devices acting in parallel. Additionally, any memory in
a computer device referred to herein may also refer to one or more memories n the
memories may be in one computing device or a plurality of computing devices acting in
parallel.
W0 2018/140984 PCT/U82018/016053
C. Operation Unit
The operation unit 214 may be configured to enable a user to interface (e.g.,
visual, audio, touch, button presses, stylus taps, etc.) with the controller 212 to l the
operation of the system 200. In some embodiments, the operation unit 214 may be further
coupled to each sensor head 204 to control the ion of each sensor head 204.
D. Display Unit
Referring again to the system 200 may further include a display unit
216 red to enable a user to view data and l information of the system 200.
The display unit 216 may further be coupled to other components of the system 200 such as
the sensor head 204. The display unit 216 may include a visual display such as a cathode
ray tube (CRT) display, liquid crystal y (LCD), light emitting diode (LED) display,
or “electronic ink” display. In some embodiments, the display unit 216 may be configured
to present a graphical user interface (e. g., a web browser and/or a client application) to the
user. A graphical user interface may include, for example, an display for GFR values as
described herein above as produced by the system 200, and operational data of the system
Exogenous Markers
Without being limited to any particular theory, molecules which are highly
hydrophilic and small (creatinine, molecular weight = 113) to moderately sized (inulin,
molecular weight ~5500) are known to be rapidly cleared from systemic circulation by
glomerular filtration. In addition to these properties, an ideal GFR agent would not be
reabsorbed nor secreted by the renal tubule, would exhibit ible binding to plasma
ns, and would have very low toxicity. In order to design optical probes that satisfy all
of these ements a balance was struck between photophysical properties, and the
molecular size and hydrophilicity of the fluorophore. For e, while hydrophobic
cyanine and indocyanine dyes absorb and emit optimally within the near infrared (NIR)
biological window (700-900 nm), hilicity is not sufficiently high to function as pure
GFR agents. Smaller dye molecules may be more easily ted to the extremely
hydrophilic species required for renal clearance, but the limited at-systems resulting from
W0 40984 PCT/U82018/016053
these lower molecular weight compounds generally enable one photon tion and
emission in the ultraviolet (UV).
To e the pharmacokinetic issues in concert with enhancing the
photophysical properties, simple derivatives of 2,5-diaminopyrazine-3,6- dicarboxylic acid
act as very low molecular weight fluorescent scaffold s with bright emission in the
yellow-to-red region of the electromagnetic spectrum. SAR studies have been carried out
using amide-linked variants of these derivatives for the simultaneous optimization of GFR
pharmacokinetics and hysical properties. A variety of hilic functionalities for
enabling rapid renal clearance of this class of pyrazine fluorophores including
carbohydrate, alcohol, amino acid and s PEG-based linker strategies may be
employed. PEG substitution maybe used to increase hydrophilicity and solubility, reduce
toxicity, and modulate ation of the resulting pyrazine derivatives. Variations of
molecular weight and architecture (and hence hydrodynamic volume) in a series of
moderately sized PEG-pyrazine derivatives may also be suitable for use as endogenous
fluorescent agents.
In one aspect, the ous fluorescent agent is lVfl3-102.
EXAMPLES
The following example illustrates various aspects of the disclosed systems
and methods.
Example I .' Sensor Head with Flared Housing
is a perspective view of a sensor head 204a in another . In
this other aspect, the sensor head 204a includes a g 600a formed from an upper
housing 602a and a flared lower housing 604a. The surface area of the lower housing 604a
expands to form an enlarged bottom surface 608a. The housing 600a further includes a
cable opening 806a formed through the upper housing 602a.
is a bottom view of the sensor head 204a showing the bottom
surface 608a of the housing 600a. The bottom surface 608a may include an aperture plate
702a including one or more apertures 704a configured to transmit light between the skin of
the patient and the light sources and light detectors contained inside the housing 600. As
W0 2018/140984
illustrated in , the apertures 704a include a light delivery aperture 1002a configured
to deliver illumination produced by the first and second light sources 218/220 to tissues of
the t 202, as well as first and second detector res 1004/1006 red to
receive light from the tissues of the t 202. In one aspect, the bottom surface 608a
enables the positioning of the apertures 704a beneath a relatively large area obscured from
t light conditions by the bottom e 608a. This reduction of scattered ambient
light entering the first and second detector apertures 1004/1006 reduces noise introduced
into the light intensity measurements obtained by the first and second light detectors
222/224.
In various aspects, the bottom surface 608a of the housing 600a may be
attached the patient’s skin using a biocompatible and transparent adhesive material 610a
including, but not limited to, a clear double-sided medical grade adhesive, as illustrated in
. The arent adhesive material 610a may be positioned on the bottom surface
608a such that the adhesive material 610a covers the apertures 704a.
is an isometric view of the sensor head 204a with the upper
housing 602a and various electrical components removed to expose an inner housing 2502.
is an exploded view of the inner housing 2502 and associated electrical
components illustrated in . ing to and , the inner g
2502 is contained within the housing 600a and is mounted to the lower g 608a. The
inner housing 2502 contains a sensor mount 912 with a first detection well 908, a second
detection well 910, and a light source well 902 formed therethrough. The first light detector
222 is mounted within the first detection well 908 and the second light or 224 is
d within the second detection well 910. The first and second light sources 218/220
are mounted within the light source well 902. In an aspect, the first detection well 908,
second detection well 910, and light source well 902 of the sensor mount 912 are optically
isolated from one r to ensure that light from the light sources 218/220 does not reach
the light detectors 222/224 without coupling through the skin of the patient 202. The
separation between the two detection wells 908/910 ensures that the detected fluorescence
signal from the exogenous fluorescent agent is distinguishable from the unfiltered
excitation light, as described in detail above.
W0 40984 2018/016053
Referring to , the inner housing 2502 includes a first ion
aperture 2602, second detection re 2604, and light source aperture 2606. The sensor
mount 912 is coupled to the inner housing 2502 so that the first detection aperture 2602,
second detection aperture 2604, and light source aperture 2606 are aligned with the first
detection well 908, second detection well 910, and light source well 902 of the sensor
mount 912, respectively.
In one , optically transparent windows 2610, 2612, and 2614 are
coupled within first detection aperture 2602, second detection re 2604, and light
source aperture 2606, respectively, to seal the apertures while also providing optically
transparent conduits between the tissues and the interior of the sensor head 204a. In
addition, diffusers 2616, 2618, and 2620 are coupled over optically transparent windows
2610, 2612, and 2614, respectively. The diffusers 2616, 2618, and 2620 are provided to
spatially homogenize light delivered to the tissues by light sources 218/220 and to spatially
homogenize light detected by light detectors 222/224. In an aspect, the absorption filter
244 is coupled to the diffuser 2616. In one aspect, an optically transparent adhesive is used
to couple the absorption filter 244 is d to the er 2616.
In view of the above, it will be seen that the several advantages of the
disclosure are achieved and other advantageous results attained. As s changes could
be made in the above s and systems without departing from the scope of the
sure, 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.
When introducing elements of the present disclosure or the various versions,
embodiment(s) or aspects thereof, the articles (4 77 (L
a an”, “the” and “said” are intended to
77 (L
mean that there are one or more of the elements. The terms “comprising
7 including” and
“having” are intended to be inclusive and mean that there may be additional elements other
than the listed elements.
Claims (12)
1. A method of monitoring a time-varying fluorescence emitted from a fluorescent agent from within a diffuse reflecting medium with time-varying l ties, the method comprising: providing a measurement data set comprising a plurality of measurement entries, each measurement data entry comprising at least two measurements obtained at one data acquisition time from a patient before and after administration of the fluorescent agent, the at least two measurements comprising an = ∗−(1: signal detected at a third region adjacent to the diffuse ting medium by a filtered light or during illumination of the diffuse reflecting medium by excitatory-wavelength light from the first region, and at least one DR signal selected from: a = signal detected at a second region adjacent to the diffuse ting ∗−(1 medium by an ered light detector during illumination of the diffuse reflecting medium by excitatory-wavelength light from a first region adjacent to the diffuse reflecting medium; a DR em signal detected at the second region by the unfiltered light or during illumination of the diffuse reflecting medium by emission-wavelength light from the first position; and a DR em,filtered signal detected at the third region by the filtered light detector during illumination of the diffuse reflecting medium by emission-wavelength light from the first on; and; identifying a gent-administration portion of the measurement data set; transforming each = ∗−(1: signal of each measurement data entry within the post-agentadministration portion of the measurement data set to an IF agent signal representing a detected fluorescence intensity emitted solely by the fluorescent agent from within the diffuse reflecting medium, wherein the transforming comprises at least one of removing the effects of leak-through of excitation-level light into the = ∗−(1: signal and removing the autofluorescence contribution to the = ∗−(1: signal; and monitoring the IF agent signal for each measurement data entry within the post-agent- administration n of the measurement data set.
2. The method of claim 1, wherein removing the effects of hrough of excitationlevel light into the = ∗−(1: signal comprises transforming each = signal into an ExLT ∗−(1 signal representing a level of excitation-wavelength light leak-through using Eqn. (21): : = = =−( ∗= = Eqn. (21) ∗−(1 where = =−( is a calibration factor.
3. The method of claim 2, wherein removing the effects of leak-through of excitationlevel light into the = ∗−(1: signal further comprises transforming each = ∗−(1: signal into a corrected fluorescence signal = ∗= ∗ =, representing detected emission-wavelength fluorescence only, using Eqn. (23): = ∗= ∗ = = = ∗−(1: −= ∗− Eqn. (23).
4. The method of claim 3, wherein ng the effects of autofluorescence comprises determining IF auto , enting intrinsic autofluorescence emitted by chromophores within the diffuse ting medium in addition to the fluorescent agent, by analyzing the = ∗= ∗ = signals obtained according to Eqn. (32): === −− = (1:= ∗(= ∗= ∗ =(1:(=(=(= ∗== ((== Eqn. (32), wherein (1:endBackground ) represents a portion of the measurement dataset ed prior to administration of the fluorescent agent.
5. The method of claim 4, wherein removing the effects of autofluorescence further comprises subtracting IF auto from = ∗= ∗ =to obtain IF agent .
6. The method of claim 2, wherein (1=−( is obtained by: ing ements from a solid phantom, the measurements comprising: a fluorescence signal = ∗−(1: representing emission-wavelength fluorescence measured using the filtered light detector; an excitation-wavelength light signal = = measured using the non-filtered ∗−(1 light detector; and computing = =−( according to Eqn. (22): ∗=−( = = ∗= ∗− Eqn. (22). = === ∗−
7. A method of determining renal on in a patient, the method comprising: providing a measurement data set comprising a plurality of measurement s, each measurement data entry comprising at least two measurements obtained at a corresponding data acquisition time from a tissue of a patient before and after administration of an exogenous fluorescent agent, the at least two measurements sing an = ∗−(1: signal detected at a third region adjacent to the diffuse reflecting medium by a filtered light detector during illumination of the diffuse reflecting medium by excitatory-wavelength light from the first region, and at least one DR signal selected from: a = signal detected at a second region adjacent to the diffuse reflecting ∗−(1 medium by an unfiltered light detector during illumination of the diffuse reflecting medium by excitatory-wavelength light from a first region adjacent to the diffuse reflecting medium; a = ∗−(1: signal detected at a third region adjacent to the diffuse reflecting medium by a filtered light detector during illumination of the diffuse ting medium by excitatory-wavelength light from the first region; a DR em signal ed at the second region by the unfiltered light detector during nation of the diffuse reflecting medium by emission-wavelength light from the first position; and a DR tered signal detected at the third region by the filtered light detector during illumination of the diffuse reflecting medium by emission-wavelength light from the first position; and; transforming each = ∗−(1: signal of each measurement data entry within the post-agentadministration portion of the measurement data set to an IF agent signal representing a detected fluorescence ity d solely by the fluorescent agent from within the diffuse reflecting medium, wherein transforming each = ∗−(1: signal ses at least one of removing the effects of leak-through of excitation-level light into the = ∗−(1: signal and removing the autofluorescence contribution to the = ∗−(1: signal; identifying a post-equilibration portion of the measurement data set; transforming the IF agent signals corresponding to the post-equilibration portion of the measurement data set to a rate of change of the IF agent signals; and determining the renal on in the patient based on the rate of change of the IF agent signals.
8. The method of claim 7, wherein removing the effects of leak-through of excitationlevel light into the = ∗−(1: signal ses transforming each = signal into an ExLT ∗−(1 signal representing a level of excitation-wavelength light leak-through using Eqn. (21): : = = =−( ∗= = Eqn. (21) ∗−(1 where = =−( is a calibration .
9. The method of claim 8, n removing the effects of leak-through of tionlevel light into the = ∗−(1: signal further comprises transforming each = ∗−(1: signal into a corrected fluorescence signal = ∗= ∗ =, representing detected emission-wavelength fluorescence only, using Eqn. (23): = ∗= ∗ = = = ∗−( −= ∗− Eqn. (23).
10. The method of claim 9, wherein removing the effects of autofluorescence comprises determining IF auto , representing intrinsic uorescence emitted by chromophores within the diffuse reflecting medium in addition to the fluorescent agent, by analyzing the = ∗= ∗ = signals obtained using the equation according to Eqn. (32): === −− = (1:= ∗(= ∗= ∗ =(1:(=(=(= ∗== ((== Eqn. (32), wherein (1:endBackground ) represents a portion of the measurement dataset obtained prior to administration of the fluorescent agent.
11. The method of claim 10, wherein removing the effects of autofluorescence further ses subtracting IF auto from = ∗= ∗ = to obtain IF agent .
12. The method of claim 8, wherein (1=−( is obtained by: ing measurements from a solid phantom, the measurements comprising: a fluorescence signal = ∗−(1: representing emission-wavelength fluorescence measured using the filtered light detector; an excitation-wavelength light signal = measured using the non-filtered ∗−(1 light detector; and computing = =−( according to Eqn. (22): ∗=−( = = ∗= ∗− Eqn. (22). = === ∗− WO 40984 113? SUBSTITUTE SHEET (RULE 26) 2131 \ QPERATEQN DESPLAY UNET UNET 212 ..... CONTRGLLER 232 ~ UNET CQNTRQL UNiT SGNTRDL UNET 2 ~ . . 222202 2222(2) 222 . ADDETEQNAL LEGHT TEMPERATURE DETECTOR(S) SENSQR(S) _ . FERST LEGHT FERST LEGHT SECQND LEGHT SECQNQ LEGHT EETECTQR SQURCE SGURCE BETECTOR SUBSTITUTE SHEET (RULE 26) WO 40984 3131 00 mmlflfl zommwfim ZMmOJwOmexwa ZOEflQOm/w Ea“ W GE % szmfixgg N/E M M “2 EMA ND “\ng Rummw .. SLENfl AHVELESHV SUBSTITUTE SHEET (RULE 26)
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Application Number | Priority Date | Filing Date | Title |
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US201762452021P | 2017-01-30 | 2017-01-30 | |
US62/452,021 | 2017-01-30 | ||
PCT/US2018/016053 WO2018140984A1 (en) | 2017-01-30 | 2018-01-30 | Method for non-invasive monitoring of fluorescent tracer agent with background separation corrections |
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