Analysis system and method for determining hemoglobin parameters in whole blood

Abstract

Absorbance is used to measure hemoglobin and bilirubin parameters in whole blood samples.SOLUTION: System 10 includes an optical-sample module 20, a spectrometer module 100, a fiber-optic assembly 90 that optically connects optical-sample module 20 to spectrometer module 100, and a processor module 150. The optical-sample module 20 includes a light-emitting module having an LED light source, a cuvette assembly, and a calibration-light module. Processor module 150 receives and processes electrical signals from spectrometer module 100 and converts the electrical signals into output signals usable to display and report hemoglobin parameter values and / or total bilirubin parameter values for a whole blood sample.SELECTED DRAWING: Figure 1

Description

Detailed Description of the Invention

[0001] [Technical field] The present invention relates generally to spectroscopic systems and methods for identifying and characterizing hemoglobin parameters in blood. [Background technology] UV-Vis spectroscopy systems include absorption spectroscopy or reflectance spectroscopy. As the name suggests, these systems use light in the visible and near-UV range to analyze samples. The wavelength range is typically from about 400 nm to about 700 nm. The absorption or reflection of visible light directly affects the perceived color of the chemicals involved. UV / Vis spectroscopy is routinely used in analytical chemistry for the quantitative determination of different analytes, such as transition metal ions, highly conjugated organic compounds, and biopolymers. Spectroscopic analysis is generally performed in solution, but solids and gases can be investigated as well.

[0002] Near-infrared spectroscopy systems also include absorption spectroscopy or reflectance spectroscopy. These systems use light in the near-infrared range to analyze samples. The wavelength range is typically from about 700 nm to less than 2,500 nm. Typical applications include pharmaceuticals, medical diagnostics (including blood glucose and pulse oximetry), food and pesticide quality control, and combustion research, as well as research in functional brain imaging, sports medicine and science, elite sports training, ergonomics, rehabilitation, neonatology, brain-computer interfaces, urology (bladder contraction), and neurology (neurovascular coupling).

[0003] Instruments for near-IR (NIR) spectroscopy are similar to those for the UV-visible and mid-IR ranges. The basic components of a spectrometer are a light source, a holder for the sample, a diffraction grating within a monochromator and prism to separate light of different wavelengths, and a detector. The radiation source is often a tungsten filament (300-2500 nm), a deuterium arc lamp (continuous across the ultraviolet range) (190-400 nm), a xenon arc lamp (continuous from 160 nm to 2000 nm), or more recently, a light-emitting diode (LED) for visible wavelengths. The detector is usually a photomultiplier tube, photodiode, photodiode array, or charge-coupled device (CCD). Single photodiode detectors and photomultiplier tubes are used in conjunction with a scanning monochromator, which filters light so that only light of a single wavelength reaches the detector at a time. The scanning monochromator moves a diffraction grating to "step through" each wavelength so that its intensity is measured as a function of wavelength. Fixed monochromators are used in conjunction with CCD and photodiode arrays. Both of these devices consist of many detectors grouped in one- or two-dimensional arrays, allowing different wavelengths of light to be collected simultaneously on different pixels or groups of pixels. Common incandescent or quartz halogen light bulbs are most often used as broadband sources of near-infrared radiation for analytical applications. Light-emitting diodes (LEDs) are also used. The type of detector used depends primarily on the range of wavelengths to be measured.

[0004] The main application of NIR spectroscopy to the human body is that the transmission and absorption of NIR light in human tissues contains information about hemoglobin concentration changes. By using several wavelength- and time-resolved (frequency or time domain) and / or spatially resolved methods, blood flow, volume, and absolute tissue saturation (StO2 or Tissue Saturation Index (TSI)) can be quantified. Applications of oximetry with NIRS methods are in neuroscience, ergonomics, rehabilitation, brain-computer interfaces, urology, and diseases affecting blood circulation (e.g., peripheral blood These include the detection of pulmonary tube diseases, the detection and evaluation of breast tumors, and the optimization of training in sports medicine.

[0005] With respect to absorption spectroscopy, the Beer-Lambert law states that the absorption of a solution is directly proportional to the concentration and path length of the absorbing species within the solution. Therefore, for a fixed path length, UV / Vis and NIR spectroscopy can be used to determine the concentration of an absorber within a solution. The method is most often used quantitatively to determine the concentration of an absorbing species within a solution using the Beer-Lambert law: A=log 10 (I0 / I)=εcL where A is the measured absorbance in absorbance units (AU), I0 is the intensity of the incident light at a given wavelength, I is the transmitted intensity, L is the path length through the sample, c is the concentration of the absorbing species.

[0006] For each species and wavelength, ε is a constant known as the molar absorptivity or extinction coefficient. This constant is a fundamental molecular property in a given solvent at a particular temperature and pressure, and is given by 1 / M * cm or often AU / M * It has units of cm. Absorbance and extinction ε are sometimes specified by natural logarithms instead of base 10 logarithms.

[0007] The Beer-Lambert law is useful for characterizing many compounds, but it is not valid as a universal relationship for concentration and absorption of all substances. Those skilled in the art will recognize that a variety of factors affect these spectroscopic systems, including spectral bandwidth, wavelength error, stray light, deviations from the Beer-Lambert law, and sources of measurement uncertainty.

[0008] Stray light is a significant factor affecting spectroscopic systems, causing the instrument to report falsely low absorbance values. Deviations from the Beer-Lambert law occur based on concentration. At sufficiently high concentrations, the absorption band will saturate and exhibit absorption flattening. The absorption peak appears flattened because close to 100% of the light has already been absorbed. The concentration at which this occurs depends on the specific compound being measured.

[0009] Measurement uncertainty arises in quantitative chemical analysis, and the results are further influenced by sources of uncertainty from the nature of the compounds and / or solutions being measured. These include absorption band overlap, fading of absorbing species (caused by decomposition or reaction), and possible compositional mismatch between sample and calibration solutions. [Summary of the Invention] Human hemoglobin (HGB) is known to be the oxygen-carrying protein in red blood cells. Determining its concentration in whole blood is a useful and important diagnostic tool in clinical biochemistry. COOx analyzers are used to measure blood hemoglobin parameters, such as total hemoglobin (tHb), carboxyhemoglobin (COHb), deoxyhemoglobin (HHb), oxyhemoglobin (OHb), methemoglobin (MeHb), and fetal hemoglobin (FHb), as well as total bilirubin (tBil), using absorbance measurements. In practice, typical COOx analyzers use lysed blood instead of whole blood due to the problems encountered in the spectroscopic analysis of whole blood. Measuring lysed blood is relatively simple. The rationale is as follows: This is because the lysis process lyses the red blood cells, making blood a nearly non-diffusive medium. The absorbance is measured by a simple collimated beam through the cuvette, where little light is lost due to scattering. Because light loss due to scattering is small, simple linear analysis can be used to find the hemoglobin and total bilirubin parameters.

[0010] Measurement of hemoglobin and total bilirubin parameters using whole blood samples is very challenging due to the strong optical scattering of whole blood. These challenges are primarily related to dealing with the increased light scattering levels of whole blood compared to lysed blood, which introduces light loss and nonlinear absorbance into the measurements.

[0011] Components within a prism-based spectrometer naturally have a low stray light profile. The primary contributor to stray light is related to how the components are used.

[0012] While the problem primarily relates to addressing the increased light scattering levels of whole blood, it is not a single factor that, if solved, can solve these challenging problems. The inventors identified several factors that need to be addressed to measure hemoglobin parameters in whole blood. Because whole blood is a highly diffusive medium, it is necessary to collect as much light as possible to reduce the requirements for the upper absorbance measurement range. Because the range of detector linearity correction is low, it is also necessary to extend the upper limit of measured absorbance. Blood sedimentation effects are another issue that result in poor correlation of absorbance in whole blood scans to absorbance in lysed blood scans. Essentially, blood cells form clumps or rouleaux. LED white light source brightness must be increased as well. Finally, new algorithms other than linear-based algorithms are needed to overcome the light scattering effects of whole blood.

[0013] Typical collection optics for systems using lysed blood are designed to collect light from the cuvette within a cone approximately ±0.7 degrees wide and have a measurement absorbance upper limit of 1.5 AU (absorbance units). For whole blood, the inventors discovered that the system needed to collect light from the cuvette within a cone approximately ±12 degrees wide, and that the upper absorbance limit had to be increased to approximately 3.5 AU. Due to blood sedimentation effects, over the typical time it takes to measure an absorbance spectrum (approximately 1 minute), the whole blood in the cuvette settles, and blood cells form clumps or rouleaux. As a result, scattering effects and absorbance change over time. The inventors discovered that changing the spectrometer controls to frequently collect multiple scans, rather than a few scans averaged over a long period of time, avoided the step function in composite absorbance scans stitched together from scans from several integration times. Unfortunately, adding more scans to extend the upper absorbance limit increases data collection time. To solve this dilemma, the integration time was reduced from 5 msec to 1.2 msec to reduce data collection time. However, it was discovered that this only worked if the light level was increased by a corresponding factor. Therefore, the LED white light brightness had to be increased.

[0014] Measuring the absorbance of diffuse samples such as whole blood presents unique challenges. The diffuse transmission of a whole blood sample perturbs the initial spatial light distribution of the measurement system, which is caused by the inherent non-uniformity of the light source. As a result, the spatial light distribution of a "blank" scan can be quite different from that of a whole blood sample scan. Because optical detectors have spatially varying responses, the response can vary with changes in the spatial distribution of the incident light, even if the overall intensity remains unchanged. An absorbance scan based on the ratio of the whole blood sample scan to the blank scan will have a significant absorbance component due to this non-uniformity of the light source in addition to the absorbance due to the sample alone. This results in significant measurement error in the whole blood sample absorbance, which is unacceptable for co-oximetry.

[0015] It was discovered that by placing the sample cuvette between diffusers, the spatial light distribution appears the same for the blank and sample scans, thus eliminating this error effect. The diffusers are specifically chosen to diffuse the incoming light rays so that they fall into the entire acceptance cone of the optical system, but not so much so that as much light throughput is preserved as possible while perturbing rays that traverse the field completely.

[0016] Furthermore, measuring fetal hemoglobin parameters presents additional challenges. These include the spectral acquisition time, which must be faster. Instead of the typical 12 seconds, the spectral acquisition time must be less than 5 seconds. The spectral acquisition time includes the integration time multiplied by the number of spectra added to each, and the processing time to generate one spectrum (full light, dark, or sample) that meets all the following requirements: The absolute wavelength accuracy is smaller; it must be less than +0.03 / -0.03 nm compared to +0.1 / -0.0 nm. Wavelength calibration maintenance (less than +0.06 / -0.0 nm compared to +0.1 / -0.0 nm), wavelength calibration drift (less than 0.024 nm / °C compared to 0.04 nm / °C), dark current levels (less than 0.06% / °C maximum dynamic range compared to 0.1% / °C maximum dynamic range), response nonlinearity (less than 0.06% after correction and less than 1.2% for the lowest and highest 10% dynamic range compared to 0.1% after correction and 2.0% for the lowest and highest 10% dynamic range), scattered light levels (less than 0.02% maximum dynamic range for the fully illuminated detector array compared to 0.1% maximum dynamic range for the fully illuminated detector array), thermal drift of response (6% intensity change maximum over the spectral range and 6% slope maximum compared to 10% intensity change maximum over the spectral range and 10% slope maximum), and temperature excursion allowed during measurements (less than 0.5°C compared to 2°C) must all be small. The present invention includes these additional features for use in measuring fetal hemoglobin parameters.

[0017] In another aspect of the present invention, commercially available compact and low-cost spectrometers typically use diffraction gratings (either reflective or transmissive) to disperse optical inputs. Diffraction gratings provide a high degree of dispersion within a small volume, producing a relatively constant bandwidth (or resolution) for wavelengths preferred by typical users. However, gratings suffer from high stray light due to multiple diffraction orders, as well as inherent imperfections in the lines etched to create the diffractive surface. Therefore, mass-produced but expensive master holographic gratings are typically used in applications requiring low stray light, rather than more commonly available replica gratings.

[0018] The low stray light requirement for COOx analyzers limits the population of suitable grating manufacturers to a few manufacturers that produce master holographic gratings or individually precision-etched gratings, which serves to make it difficult to obtain low-cost, high-performance gratings in volume.

[0019] Prisms are also used to make spectrometers. Prisms do not have any issues with multiple diffraction orders, and their surfaces have orders of magnitude less imperfections than those of gratings. Components within a prism-based spectrometer naturally have a low stray light profile. Therefore, stray light in a prism spectrometer can be an order of magnitude or more lower compared to a grating spectrometer of otherwise similar design. The major contributors to stray light performance arise from how the components are used. There are three main sources of stray light. These include (1) overfilling of the spectrometer numerical aperture, (2) retroreflections from the optical array detector, and (3) focal plane images. Light in excess of that required to fully illuminate the spectrometer numerical aperture can bounce around within the spectrometer and land on the detector. In the present invention, optical filters The numerical aperture of the fiber is 0.22, and the numerical aperture of the prism spectrometer is 0.1. An aperture placed above the optical fiber input limits the light input cone from the optical fiber to prevent excessive light input. The optical array detector does not absorb all of the light incident on it, but rather reflects a portion back. This retroreflection must be controlled to land within an absorbing surface or beam trap to prevent scattering on the detector. A slight tilt of the optical array detector forces the retroreflection back in a harmless direction. The image of the slit on the detector focal plane should be as sharp as possible. Any excessive overfill of the detector due to defocus can be a potential stray light source. If this light strikes detector structures such as bond wires, metallization pads, etc., it can bounce back onto the sensitive surface of the detector.

[0020] Furthermore, prism spectrometers spread the bluish color of the spectrum over more pixels than grating spectrometers, and therefore, the bluish color of the spectrum gives a lower signal for the pixel. To compensate for the lower signal for the pixel, an LED with higher blue power or a bluish white LED is used. The blue signal can be further increased by adding an inexpensive filter glass after the LED that slightly attenuates the reddish color. Kopp filter glass type 4309, approximately 3 mm thick, is useful for this purpose. The main disadvantages of prisms are their lower dispersion power compared to gratings and the variation in resolution with wavelength. In the present invention, in which prisms are used, the former disadvantage is mitigated by using a sufficiently small optical array detector, and the latter disadvantage is mitigated because the analysis of whole blood does not require uniformly small resolution across the waveband of interest.

[0021] Currently available spectrometers typically provide a uniform 1 nm resolution for the blood measurement spectral range of 455 to 660 nm. In the present invention, the spectral range is expanded to cover the 422 to 695 nm spectral range. Furthermore, the resolution is selectively increased in regions where lower resolution is not required (such as the 600 to 695 nm and 422 to 455 nm regions). In the present invention, these regions have a resolution greater than 1 nm. Typically, the resolution is about 3.0 nm to about 3.5 nm. These ranges are used to capture additional wavelength calibration peaks for wavelength calibration and fluid detection. The larger spectral range of the present invention requires consideration of the dispersion spectrum from the prism. The dispersion spectrum must be spread across the optical array detector, not covering enough pixels to extend beyond the detector array, but covering enough pixels to sample the spectrum with sufficiently fine resolution. Due to the wider spectral range, the present invention incorporates an optical array detector with 1024 pixels and an active area length of about 8.0 mm.

[0022] A minimal partial reference design for an optically dispersive spectrometer requires only two components: an optically dispersive element (i.e., a prism or grating) and a doublet (achromatic) lens. The prism / grating has a reflective coating on its base. An example of an acceptable prism is a Littrow prism. A Littrow prism has a structure such that it is useful for the compact, low-cost spectrometer of the present invention. The prism material (dispersion characteristics) and lens focal length are further considerations. Other While prisms and achromatic lenses may be used, one embodiment of the present invention incorporates a Schott F5 glass prism and an 80 mm focal length lens. This particular combination provides a spectral dispersion length of approximately 6.48 mm. This dispersion length leaves approximately 0.75 mm on either side of the optical array detector available for tolerance variations and dark correction pixels.

[0023] Thermal drift of the spectral response must be considered. It is very important that the spectral response of the spectrometer stays within a certain range between full light and whole blood scans. Any change in the spectrometer will cause absorbance errors. The primary prevention against this change is The solution is to ensure that the image of the slit overfills the pixel so that thermal image drift does not cause a reduction in light on the detector pixel. 1:1 imaging of the system combined with a 200 μm diameter optical fiber overfills a 125 μm tall pixel. Thermal drift is not an issue as long as the image drift is limited to less than about 30 μm of movement in either direction along the detector over the measurement interval. The present invention also contemplates various mechanisms for minimizing the effects of thermal drift on the spectral response. These mechanisms include insulating the spectrometer housing to minimize temperature changes external to the spectrometer housing, maintaining the temperature within the spectrometer housing using a temperature-controlled heat source, and / or incorporating temperature-compensated lens mounts for achromatic lenses.

[0024] The inventive process for converting the electrical signal from a spectrometer will now be discussed. First, the absorbance is measured, and the absorbance is subtracted from the base 10 logarithm of the ratio of the electrical signal received when a blood sample is in the cuvette to the electrical signal received when a clear fluid is in the cuvette. Second, the absorbance values ​​at each wavelength are placed into a mapping function, which maps the absorbance values ​​to analyte levels (COOx parameters and bilirubin) in the whole blood sample. The mapping function and its coefficients are established by using the absorbance values ​​measured for the whole blood sample along with known analyte values ​​and establishing a relationship between these absorbance values ​​and the known analyte values.

[0025] The present invention achieves these and other objectives by providing a compact, low-cost COOx analyzer subsystem. In one embodiment of the present invention, there is a system for measuring hemoglobin parameters of a whole blood sample, the system including: (a) an optical sample module having a light emitting module, a replaceable cuvette assembly, and a calibration light module; (b) an optical fiber; (c) a spectrometer module; and (d) a processor module. The light emitting module has an LED light source capable of emitting light, the light being directed along an optical path. The cuvette assembly is adjacent to the light emitting module, and the cuvette assembly is adapted to receive a whole blood sample and has a sample receiving chamber having first and second cuvette windows aligned with each other. The sample receiving chamber is disposed in the optical path to receive light from the LED light source and has a defined optical path length between the first and second cuvette windows, along with an electronic chip capable of storing a path length value of the sample receiving chamber. The calibration light module has a calibration light source having light of one or more known wavelengths, and the calibration light module is capable of emitting the calibration light into the optical path. The optical fiber has a light receiving end and a light emitting end. The receiving end is optically connected to the optical sample module, which receives light from the optical path and transmits the light to the emitting end. The spectrometer module receives light from the emitting end of the optical fiber, separates the light into multiple light beams, each having a different wavelength, and converts the multiple light beams into electrical signals. The processor module obtains (1) a value for the path length of the sample receiving chamber from the electronic chip and processes (2) the electrical signal generated from the spectrometer module for the whole blood sample. The value for the path length of the sample chamber is used to convert the electrical signal into an output signal usable for displaying and reporting a hemoglobin parameter value and / or a total bilirubin parameter value for the whole blood sample.

[0026] In another embodiment of the present invention, the light emitting module includes a plurality of optical components disposed in an optical path between the LED light source and the cuvette assembly, the plurality of optical components including at least one optical diffuser and one or more of a collimating lens, a circular polarizer, and a converging lens. In a further embodiment of the present invention, the calibration light module includes a diffuser plate disposed in the light path downstream from the cuvette assembly but upstream from the beam splitter. In yet another embodiment of the present invention, an absorbance measurement system for whole blood is disclosed. The system includes an optical sample module, an optical fiber, a spectrometer module, and a processor module. The optical sample module includes a light emitting module, a cuvette module, a first optical diffuser, and a second optical diffuser. The cuvette module is positioned between the first optical diffuser and the second optical diffuser. The spectrometer module receives light from the light emitting end of the optical fiber, separates the light into multiple light beams, and converts the multiple light beams into electrical signals. The processor module receives and processes electrical signals generated for the whole blood sample from the spectrometer module and converts the electrical signals into output signals usable for displaying and reporting hemoglobin and / or total bilirubin parameter values ​​for the whole blood sample. In yet another embodiment, the spectrometer module includes an input slit positioned in the optical path to receive light emitted from the emitting end of the optical fiber and transmit the light through the input slit; an optically dispersive element disposed in the optical path to receive the light transmitted through the input slit, separate the light into multiple light beams each having a different wavelength, and redirect the multiple light beams toward but offset from the input slit; and an optical array detector capable of receiving the multiple light beams and converting the multiple light beams into electrical signals for further processing. In another embodiment, the spectrometer module includes thermal compensation means for maintaining the positions of the plurality of light beams on the optical array detector, the thermal compensation means including one or more of a thermal insulator disposed around the spectrometer housing, a temperature controller assembly disposed on the spectrometer housing (e.g., a heating tape having a thermistor or other temperature measuring component and a program that controls heating of the tape based on the temperature within the spectrometer housing), and a thermal compensation lens mount. In a further embodiment, the thermal compensation lens mount has a fixed mount end and a loose mount end that allows for thermal expansion and contraction of the thermal compensation lens mount. The fixed mount end is fixedly attached to the bottom of a base plate or spectrometer housing. The lens mount has an expansion coefficient greater than the expansion coefficient of the base plate or spectrometer housing to which the lens mount is attached. The thermal compensation lens mount moves linearly and laterally relative to the optical path of the light from the optical input slit based on the expansion coefficient of the lens mount. This temperature-based movement of the lens mount maintains the position of the dispersed light from the optical dispersive element on the optical array detector. In other words, thermal repositioning of the achromatic lens by the thermal compensation lens causes the dispersed light from the dispersive element to be incident on the optical array detector without affecting the electrical signal generated by the optical array detector from the incident light. The shift of the light beam is caused by the optical dispersive element, which responds to temperature changes. In another embodiment, a compact spectrometer for measuring hemoglobin parameters in whole blood is disclosed. The spectrometer includes an enclosed spectrometer housing having an optical fiber receiving end with an optical input end / light incidence port, an optical input slit disposed on a circuit board substrate disposed within the enclosed spectrometer housing, the optical input slit aligned with and adjacent to the optical input port, an optical array detector disposed on the circuit board substrate adjacent to the optical input slit, and an optical component group consisting of an optical dispersive element disposed downstream from the optical input slit and a spherical achromatic lens disposed between the optical input slit and the optical dispersive element, the optical dispersive element having a reflective surface on its back surface for reflecting dispersed light back toward the achromatic lens. The achromatic lens transmits light from the optical input slit to the optical dispersive element and transmits the dispersed light reflected from the optical dispersive element to the optical array detector. To achieve this, the achromatic lens is slightly off-axis with respect to the light coming from the light input slit, so that the dispersed light from the light dispersive element is directed back to the light array detector rather than back to the light input slit. In a further embodiment, a method for measuring whole blood hemoglobin parameters despite strong optical scattering caused by whole blood is disclosed. The method includes providing a light source, such as an LED light source, having a spectral range of about 422 nm to about 695 nm; directing light having the spectral range from the light source along an optical path; a cuvette module having a sample-receiving chamber with a first cuvette window disposed in the optical path, the first cuvette window configured to receive the sample; providing a cuvette module that transmits light to a sample receiving chamber of the cuvette and a second cuvette window aligned with the first cuvette window, the sample receiving chamber containing a whole blood sample; providing a pair of diffuser plates (i.e., a first diffuser plate and a second diffuser plate) disposed in the light path, wherein the first cuvette window of the sample receiving chamber of the cuvette and the second cuvette window are disposed between the pair of diffuser plates; directing the light from the cuvette module into a spectrometer having a light-dispersing element that separates the light into multiple light beams each having a different wavelength and converts the multiple light beams into electrical signals; and processing the electrical signals into output signals usable for displaying and reporting a hemoglobin parameter value and / or a total bilirubin parameter value of the whole blood sample. In another embodiment of the method, the processing step includes processing the electrical signal into spectral absorptances and then mapping the spectral absorptances to hemoglobin and / or bilirubin parameter values ​​using a computational mapping function. In yet another embodiment of the method, the processing step includes using a kernel-based orthogonal projection to latent structure mapping function as the calculated mapping function.

[0027] In yet another embodiment of the present method, a method for measuring hemoglobin parameters in a whole blood sample is disclosed, the method including: (1) measuring and recording transmitted light intensity scans across multiple wavelengths within a measurement range by transmitting light through a cuvette module, the cuvette module having an optical path with a known optical path length through the cuvette module and filled with a transparent fluid; and (2) measuring and recording transmitted light intensity scans across multiple wavelengths within the measurement range by transmitting light a second time through a cuvette, the cuvette having an optical path with a known optical path length through the cuvette and filled with a whole blood sample, wherein the steps of measuring and recording the transmitted light for the transparent fluid and the whole blood sample, respectively, include: (3) determining a spectral absorptance at each of a plurality of wavelengths in a measurement range based on a ratio of a transmitted light intensity scan of the whole blood sample to a transmitted light intensity scan of a transparent fluid using a prism-based spectrometer; and (4) correlating the absorptance at each of a plurality of wavelengths in the measurement range with a hemoglobin parameter value and / or a bilirubin parameter value of the blood sample using a computational mapping function. [Brief explanation of the drawings]

[0028] [Figure 1] FIG. 1 is a simplified perspective view of one embodiment of the present invention showing a compact COOx subsystem. [Figure 2] FIG. 2 is a side elevation view of one embodiment of the optical sampling module shown in FIG. 1. [Figure 3] 3 is a front perspective view of one embodiment of the light emitting module of the optical sample module shown in FIG. 2. FIG. [Figure 3A] 4 is a front perspective view of one embodiment of the light emitting module shown in FIG. 3, illustrating several optical components. [Figure 3B] FIG. 3B is an enlarged side elevational view of the optical component shown in FIG. 3A. [Figure 4]2 is a front perspective view of one embodiment of a cuvette assembly of the optical sample module shown in FIG. 1. [Figure 5] FIG. 5 is a rear perspective view of the cuvette assembly shown in FIG. 4. [Figure 6] FIG. 1 is a front elevation view of a cuvette module of a cuvette assembly, showing the fluid input and output ports, the sample receiving chamber, the sample window, and the electronic chip assembly. [Figure 7] 7 is a rear perspective view of the sample-receiving chamber of FIG. 6, showing the cuvette first and second windows. [Figure 8] FIG. 2 is a rear plan view of the sample-receiving chamber showing the electronic chip assembly disposed adjacent to the sample-receiving chamber. [Figure 9] 2 is a perspective view of one embodiment of a calibration light module of the optical sample module of FIG. 1. [Figure 10] 9 is a cross-sectional side view of the calibration light module of FIG. 8, showing the calibration light source. [Figure 11] 10 is a schematic side plan view of the calibration light source of the calibration light module of FIG. 9, showing several optical components. [Figure 12] FIG. 2 is a front perspective view of one embodiment of the spectrometer module of FIG. 1 with the cover removed to show the internal components. [Figure 13] FIG. 13 is a rear perspective view of the spectrometer module of FIG. 12 showing the input light slit and adjacent light array detector. [Figure 14] FIG. 13 is a cross-sectional rear view of the spectrometer module of FIG. 12 showing the single circuit board and the location of the input light slit and light array detector. [Figure 15] FIG. 13 is a top view of the spectrometer module of FIG. 12 showing the optical components with superimposed ray traces. [Figure 16] 1 is a ray trace showing input light from an input light slit and multiple light beams refracted onto a light array detector. [Figure 17A] FIG. 1 is a perspective view of one embodiment of a thermal compensation means for a spectrometer module, showing insulation wrapped around the spectrometer module. [Figure 17B]FIG. 10 is a perspective view of another embodiment of a thermal compensation means for a spectrometer module, showing a temperature control assembly. [Figure 17C] 13 is a cross-sectional view of one embodiment of a lens mount of the spectrometer module of FIG. 12, showing a temperature compensated lens mount. [Figure 18] 13 is a cross-sectional view of one embodiment of a lens mount of the spectrometer module of FIG. 12, showing a fixed lens mount. [Figure 19] FIG. 12 is a graphical representation showing correlation results of the COOx analyzer subsystem of the present invention for total hemoglobin using the K-OPLS mapping function and method. [Figure 20] FIG. 10 is a graphical representation of correlation results of the COOx analyzer subsystem of the present invention for oxygenated hemoglobin using the K-OPLS mapping function and method. [Figure 21] FIG. 1 is a graphical representation showing correlation results of the COOx analyzer subsystem of the present invention for carboxyhemoglobin using the K-OPLS mapping function and method. [Figure 22] FIG. 10 is a graphical representation of correlation results of the COOx analyzer subsystem of the present invention for deoxygenated hemoglobin using the K-OPLS mapping function and method. [Figure 23] FIG. 12 is a graphical representation showing correlation results of the COOx analyzer subsystem of the present invention for methemoglobin using the K-OPLS mapping function and method. [Figure 24] FIG. 1 is a graphical representation showing correlation results of the COOx analyzer subsystem of the present invention for total bilirubin using the K-OPLS mapping function and method. DETAILED DESCRIPTION OF THE INVENTION

[0029] [Detailed Description of the Invention] Embodiments of the present invention are illustrated in Figures 1-24. Figure 1 illustrates one embodiment of a COOx analyzer subsystem 10. The COOx analyzer subsystem 10 includes at least an optical sample module 20, an optical fiber 90, and a spectrometer module 100. The COOx analyzer subsystem 10 can optionally include a processor module 150, or the processor module 150 can optionally be included in the electronics of a diagnostic system of which the COOx analyzer subsystem 10 is a part. Line 5 is included to indicate that the processor module 150 may or may not be part of the COOx analyzer subsystem 10. The processor module 150 includes, but is not limited to, a microprocessor module 152 and a memory module 154. Optionally, the processor module 150 also includes a converter module 156. The converter module 156 can be included in the COOx analyzer subsystem 10, or can be external to the COOx analyzer subsystem 10. The COOx analyzer subsystem 10 uses absorbance to It is used to measure blood hemoglobin parameters such as total hemoglobin (tHb), carboxyhemoglobin (COHb), deoxyhemoglobin (HHb), oxygenated hemoglobin (OHb), methemoglobin (MetHb), and fetal hemoglobin (FHb), as well as total bilirubin (tBil).

[0030] 2 shows optical sample module 20. Optical sample module 20 includes light emitting module 22, cuvette assembly 40, and calibration light module 60. Light emitting module 22, as the term suggests, emits a visible light beam toward cuvette assembly 40, which is then received by calibration light module 60, which then transmits the visible light beam to spectrometer module 100. Light beam 12 defines optical path 21.

[0031] 3-3A show perspective views of an embodiment of light emitting module 22 of FIG. 2. Light emitting module 22 includes a light emitting module substrate 24 that includes electrical circuitry (not shown) and a light emitting optics assembly 25. Light emitting optics assembly 25 has an optics assembly housing 26 having an optics assembly end 26a. A visible light beam 28a emerges from optics assembly end 26a of light emitting optics assembly 25 when light emitting module 22 is powered on by a signal received from processor 150. FIG. 3A shows light emitting optics assembly 25 with optics assembly housing 26 removed, exposing multiple optical components B contained within light emitting optics assembly 25.

[0032] Referring now to FIG. 3B, an enlarged side view of several optical components B from FIG. 3A is shown. In this embodiment, component B includes a light-emitting diode (LED) light source 28, a collimating lens 30, a first diffuser 32, a circular polarizer 34, a converging lens 36, and an optional protective window 38. The circular polarizer 34 offers a particular advantage: improved sensitivity and accuracy of the system. Hemoglobin has optical rotation properties, which mean that if non-circularly polarized light is used to measure hemoglobin absorbance, the polarization sensitivity of the spectrometer will cause absorbance errors. Unlike other polarization states of light, the polarization state of circularly polarized light does not change when passing through hemoglobin. Therefore, the polarization response of the spectrometer is the same when circularly polarized light passes through hemoglobin, as is the case for a reference scan performed on a cuvette filled with a transparent fluid.

[0033] 4 and 5 show front and back perspective views of one embodiment of a cuvette assembly 40. The cuvette assembly 40 includes a cuvette base 41 and a cuvette module 43. The cuvette base 41 provides support for fastening the cuvette assembly 40 within the analyte subsystem 10 and includes a cuvette light path opening 42 that is disposed within the light path 21 and aligned with the light beam emitted from the light emitting module 22. The cuvette module 43 includes a cuvette first portion 44 having a sample receiving recess 45, a sample inlet port 46, a sample outlet port 47, an electronic chip assembly 48, and a first cuvette window 49, and a cuvette second portion 50 having a second cuvette window 52 (shown in FIG. 6 and outlined as line 53) opposite and aligned with the first cuvette window 49, such that the first and second cuvette windows 49, 52 are aligned with and dispersed within the optical path 21. The cuvette first portion 44 and the cuvette second portion 50 are bonded together with or without a gasket disposed between the cuvette first portion 44 and the cuvette second portion 50. Bonding can be achieved using adhesives, ultrasonic techniques, solvent-based techniques, or the like. 6, the sample receiving recess 45 of the cuvette first portion 44 forms a sample receiving chamber 54 with the cuvette second portion 50 in fluid communication with the sample inlet and outlet ports 46, 47. The distance between the cuvette first portion 44 and the cuvette second portion 50 of the sample receiving chamber 54 defines the cuvette optical path length. The cuvette path length is precisely measured and stored in the electronic chip 48 for later retrieval by the processor module 150. A typical path length used in this embodiment of the invention is 0.0035 inches (0.090 mm).

[0034] Referring now to FIG. 7 , an enlarged rear perspective view of the first and second cuvette portions 44, 50 is shown. As shown, the first cuvette portion 44 has a sample chamber recess 45 with a first cuvette window 49 and an electronic chip recess 48a for receiving the electronic chip assembly 48. The second cuvette portion 50 has a second cuvette window 52 that forms a sample-receiving chamber 54 when assembled with the first cuvette portion 44. The second cuvette window 52, ​​delineated by an outline 53 on the second cuvette portion 50, is a raised surface that forms a waterproof seal around the sample chamber recess 45 and the sample-receiving chamber 54. Optionally, a thin gasket can be positioned between the first and second cuvette portions 44, 50 to more easily ensure a waterproof seal. FIG. 8 shows a rear view of the first cuvette portion 44 with the electronic chip assembly 48 disposed within the electronic chip recess 48a. The electronic chip assembly 48 includes a chip circuit board 48b and an electronic chip 48c that stores the cuvette path length value for the particular cuvette module 43. The first cuvette window 49 is disposed in the optical path 21 and transmits the light beam passing through the sample to the calibration light module 60, which then passes the light beam to the spectrometer module 100.

[0035] 9, there is shown one embodiment of a calibration light module 60. The calibration light module 60 includes a calibration module housing 62, a light beam receiving portion 64, a calibration light portion 70, and an optical fiber portion 80, with the calibration module housing 62 and the light beam receiving portion 64 aligned with the light path 21.

[0036] 10 is a cross-sectional elevation view of calibration light module 60. Calibration module housing 62 includes a first tubular conduit 62a between a light beam input aperture 62b and a light beam exit aperture 62c, as well as a calibration light beam aperture 62e at an opposite end that traverses and intersects first tubular conduit 62a.

[0037] The light beam receiving portion 64 houses a collimating lens 66 that collimates the light beam 28a received along the light path 21 from the cuvette module 43 and directs the light beam 28a into the first tubular conduit 62a. Disposed within the calibration module housing 62 is a beam splitter holder assembly 67 disposed transversely across the first tubular conduit 62a. The beam splitter holder assembly 67 has an upwardly sloping surface 67a that faces the calibration light beam opening 62e and the light beam exit opening 62c in the light path 21. The beam splitter holder assembly 67 supports a second diffuser plate 68 and a beam splitter 69 (shown in FIG. 11), the beam splitter 69 being disposed downstream from the second diffuser plate 68 along the optical path 21 so as to be positioned to receive the calibration light beam 72a and direct the calibration light beam 72a along the optical path 21 and the first tubular conduit 62a to the light beam exit aperture 62c.

[0038] Calibration light portion 70 includes a calibration light source 72 disposed adjacent to but spaced apart from optical path 21 and capable of directing a calibration light beam 72a across optical path 21 through calibration light opening 62e into calibration module housing 62 toward beamsplitter holder assembly 67. Within calibration light portion 70 is a collimating lens 74 that collimates calibration light beam 72a before being reflected by beamsplitter holder assembly 67 toward light beam exit opening 62c.

[0039] Optical fiber portion 80 is positioned in optical path 21 at or adjacent to light beam exit aperture 62c. The fiber optic portion 80 includes a fiber optic connector assembly 84 that includes a focusing lens 82 and a connector housing 80 that is adapted to receive the fiber optic assembly 90. The fiber optic portion 80 is adapted to ensure that the light beam 28a is properly focused into the fiber optic assembly 90 by the focusing lens 82.

[0040] Figure 11 is a schematic diagram of Figure 10, showing the relative positions of optical components 66, 68, 69, 74, and 82, light beams 28a and 72a, and fiber optic assembly 90. As can be seen in Figure 11, light beam 28a is received by collimating lens 66 and transmitted through second diffuser 68 and beam splitter 69 to focusing lens 82 and into fiber optic assembly 90. As discussed above, the importance of using a diffuser pair (first diffuser 32 and second diffuser 68) with cuvette module 43 between diffuser pair 32 and 68 is that the spatial light distribution will appear the same for blank scans and whole blood sample scans. The use of diffusers 32 and 68 in this arrangement eliminates error effects caused by variations in light source nonuniformity and / or variations in the spatial distribution of the incident light, even if the overall intensity remains unchanged. The diffusers 32, 68 are selected to diffuse the incident light rays into the entire acceptance cone of the optical component group 120 of the spectrometer module 100. This effectively scatters the rays completely across the optical measurement field.

[0041] When activated, the calibration light beam 72a is received by the collimating lens 74, transmitted to the beam splitter 69, directed to the converging lens 82, and focused into the optical fiber assembly 90. The calibration light beam 72a has a unique wavelength of light used to calibrate the wavelength scale of the spectrometer module 100. One example of an acceptable calibration light source 72 is a krypton (Kr) gas discharge lamp, which provides seven Kr line wavelengths in nanometers covering the range from 422 nm to 695 nm. The prism 131 of the optical dispersion component 130 has nonlinear dispersion versus wavelength, requiring a higher-order polynomial or other function. The present invention uses a fifth-order polynomial for the pixel location of the Kr line peaks, providing residuals well below the absolute wavelength accuracy requirement of + / - 0.03 nm.

[0042] The optical fiber assembly 90 includes an optical fiber 92, a first optical fiber connector 94, and a second optical fiber connector 96 (shown in FIG. 12). The first optical fiber connector 94 is clipped onto an optical receiving end 92a of the optical fiber 92 and is directly and removably connected to the connector housing 86 of the optical fiber connector assembly 84. One embodiment of the optical fiber 92 includes a 200 μm silica core fiber having a numerical aperture (NA) of 0.22.

[0043] 12 and 13, one embodiment of a spectrometer module 100 is shown. Spectrometer module 100 includes a spectrometer housing 102, a spectrometer base 104, a spectrometer cover 106 (shown in FIG. 1), an optical fiber housing end 108, and an electrical signal output coupler 103. Spectrometer module 100 has outer envelope dimensions of 11 cm x 8 cm x 2 cm and optionally includes a thermal compensation structure, discussed below. Contained within spectrometer housing 102 are the essential components of spectrometer module 100. These components include an optical receiving and converting assembly 110 and an optical component group 120. Optical component group 120 includes an achromatic lens assembly 121 and an optically dispersive element 130. Optically dispersive element 130 can be a prism 131 or a grating 136. Fiber optic assembly 90 is removably fastened to fiber optic housing end 108 at optical entrance port 109, and fiber optic assembly 90 transmits light beams 28a, 72a to spectrometer module 100. As previously mentioned, light beam 28a represents light transmitted from light emitter module 22 through cuvette module 43, while light beam 72a represents calibration light transmitted from calibration light module 60 used to calibrate spectrometer module 100.

[0044] Achromatic lens assembly 121 includes a lens mount 122 and a spherical achromatic lens 124. Achromatic lens 124 receives light beam 28a, 72a, as the case may be, and directs the light beam to an optically dispersive element 130, which in this embodiment is a prism 131. Prism 131 has a reflective coating 132 on its outer rear surface. Prism 131 refracts light beam 28a and reflects the light back through achromatic lens 124.

[0045] The optical receiving and converting assembly 110 is securely mounted adjacent the inner surface 108a of the optical fiber housing end 108. The optical receiving and converting assembly 110 includes a circuit board substrate 112 on which an optical input slit 114 is mounted, which aligns with the emitting end 92b (not shown) of the optical fiber 92. Adjacent to the input slit 114 is an optical array detector 116 that receives the refracted light from the prism 131. The optical array detector 116 converts the refracted light into an electrical signal, which is output to the processor module 150 through an output connector 118. Providing the optical input slit 114 and the optical array detector 116 adjacent to each other on the circuit board 112 has several advantages. This feature significantly simplifies the construction and improves the accuracy of the spectrometer module 100. Other spectrometers place these items on separate planes, which have separate mounting structures and must be adjusted independently. This feature of mounting the input slit and optical array detector adjacent to each other on circuit board 112 eliminates the need to separately mount and position each structure (ie, slit and detector).

[0046] FIG. 14 is an enlarged view of the optical receiving and converting assembly 110. The optical input slit 114 is 15 μm wide by 1000 μm long and projects a rectangular optical fiber slit image, approximately 15 μm wide by 200 μm high, onto the optical array detector 116 (Hamamatsu S10226-10 is an example of an optical array detector that can be used). The input slit 114 is applied directly onto the same circuit board substrate 112 as the optical array detector 116 and in close proximity thereto. The optical array detector 116 has a pixel height between approximately 100 μm and approximately 150 μm, allowing one-to-one imaging of a 200 μm diameter optical fiber onto the detector. In this embodiment, the input slit 114 is laser etched in a precise location relative to the optical array detector 116, making alignment less labor-intensive. Because the input slit 114 and the optical array detector 116 are only slightly off-axis with respect to the central axis of the achromatic lens 124, there are minimal aberrations and one-to-one imaging onto the optical array detector 116 is possible, so no cylindrical focusing lens is required to shrink the optical fiber image (200 μm diameter fiber) to match the pixel height of the optical array detector 116.

[0047] Referring now to FIG. 15, there is shown a top view of the spectrometer module 100 of FIG. 13. Superimposed on FIG. 15 is a ray trace diagram 140 of a light beam transmitted to the spectrometer module 100 by the optical fiber 92. As shown, light beam 28a enters the spectrometer module 100 and passes through the input slit 114 toward the achromatic lens 124. The achromatic lens 124 is used off-axis; that is, the achromatic lens is slightly off-axis relative to light beam 28a. Light beam 28a is transmitted by the achromatic lens 124 to a prism 131, which refracts light beam 28a, as a prism should, into multiple light beams 138a, 138b, and 138c of different wavelengths. The multiple light beams 138a, 138b, and 138c are reflected by the prism 131 back through the achromatic lens 124. An achromatic lens 124 is used off-axis to direct the multiple refracted and reflected light beams 138 a , 138 b , 138 c from the prism 131 onto the optical array detector 116 .

[0048] 16 is an enlarged view of the ray trace diagram 140. The achromatic lens 124 It is used off-axis with respect to the incoming light beam 28a. By using the achromatic lens 124 off-axis in conjunction with a prism 131 having a reflective coating 132 on its base, a compact, minimum component spectrometer module 100 is achieved that can be used to measure hemoglobin and / or total bilirubin parameters in whole blood.

[0049] Temperature changes have a significant effect on the beam refraction angle when a prism is used instead of a diffraction grating. In the present invention, a thermal compensation means 160 is provided to compensate for the thermal shift of the incoming light beam caused by the optical dispersive element 130. Temperature changes within the spectrometer module 100 cause thermally induced movement of the slit image from the input slit 114 on the optical array detector 116, caused by thermally induced changes in the refractive index of the dispersive prism 131. FIG. 16 shows, by arrow 400, the movement direction of the image on the optical array detector 116 for thermally induced refractive index changes in the prism 131. If the lens 124 were to move in the opposite direction over the same temperature interval, as indicated by arrow 402, the slit image would move back to where it should be on the optical array detector 116. To prevent this shift, thermal compensation means 160 may be as simple as wrapping spectrometer module 100 in insulation to minimize temperature changes within spectrometer module 100 from temperature changes occurring outside of spectrometer module 100, or by placing spectrometer module 100 in a temperature-controlled space. Another approach is to include a temperature controller assembly 170 that maintains a predefined constant temperature, including at least a ribbon heater 172 attached to the inside or outside surface of spectrometer housing 102, and a temperature sensor 174, such as a thermocouple or thermistor, for measuring the temperature of the spectrometer housing and heater circuitry. Figures 17A and 17B illustrate these possibilities.

[0050] In one embodiment, shown in FIG. 17C , the achromatic lens mount 122 is a thermally compensated lens mount. The thermally compensated lens mount 122 has a fixed mount end 122 a and a loose mount end 122 b. The fixed mount end 122 a is fixedly fastened to the spectrometer base 104 or a base plate 104 a that is rigidly attached to the spectrometer base 104. The loose mount end 122 b typically has a fastener 126 that extends through a lens mount slot 122 c of the lens mount 122 and into the spectrometer base 104 or base plate 104 a. A hold-down spring 128 exists between the head 126 a of the fastener 126 and the lens mount 122. Sufficient clearance exists between the lens mount slot 122 c and the fastener 126 to allow for expansion / contraction of the lens mount 122 due to temperature changes. Because the coefficient of expansion of lens mount 122 is greater than the coefficient of expansion of spectrometer base 104 and / or base plate 104a, free mount end 122b allows for thermal expansion and contraction of thermally compensated lens mount 122 in the direction indicated by arrow 500, which is linear and transverse to the light beam from input slit 114. This configuration allows the achromatic lens to slide relative to base plate 104a and / or other components mounted to spectrometer base 104. Thermally compensated lens mount 122 ensures that multiple light beams 138a, 138b, 138c will always impinge on optical array detector 116 with sufficient intensity despite temperature changes within spectrometer housing 102 without affecting the electrical signal generated by optical array detector 116. One such material that meets the requirement that the lens mount 122 (in some cases) have a larger coefficient of expansion than the spectrometer base 104 and / or base plate 104a is plastic, which is a modified polyphenylene ether (PPE) resin sold under the trademark NORYL®, which consists of an amorphous blend of polyphenylene oxide (PPO), polyphenylene ether (PPE) resin, and polystyrene.

[0051] FIG. 18 shows an alternative embodiment of the lens mount 122. In this embodiment, the lens mount 122 has two fixed mount ends 122a, each fastened to the base plate 104a and / or the spectrometer base 104 by a fastener 126. Because both ends 122a of the lens mount 122 are fixed, any temperature changes within the spectrometer module 100 will affect the angles at which the multiple light beams 138a, 138b, and 138c impinge on the optical array detector 116. As previously disclosed with respect to the slit image and the length of the optical array detector 116, a temperature change of greater than 0.5°C will cause one of the light beams of that intensity to not fully impinge on the optical array detector 116, thus resulting in an inaccurate reading. To null this possible effect, the spectrometer module 100 is equipped with a temperature controller assembly (not shown) to ensure that the prism 131 and the achromatic lens assembly 121 remain at a constant temperature. While there are several methods available for maintaining a constant temperature inside the spectrometer module 100, one example of such a temperature controller assembly for achieving this is a ribbon heater having a thermistor adhesively attached to the inside or outside of the spectrometer module 100, which is controlled by an electronic regulation circuit (not shown). Optionally, the spectrometer module 100 can also be insulated on the inside, outside, or both to more easily maintain a given temperature and to protect against temperature changes in the vicinity surrounding the spectrometer module 100. Other mechanisms include placing the spectrometer module 100 in a temperature-controlled environment.

[0052] Training data: A dataset of approximately 180 blood samples from approximately 15 different individuals was developed. Blood samples were manipulated using sodium nitrite to elevate MetHb levels and CO gas to elevate COHb levels. Plasma was removed from or added to the samples to alter tHb levels. Bilirubin spiking solutions were added to vary tBil levels. A tonometer was used to manipulate oxygen levels. Blood samples were manipulated to cover a wide range of analyte values. Blood samples were then measured on a reference lysate pHOx Ultra analyzer equipped with a COOx analyzer and analysis software. Whole blood spectra were collected on a pHOx Ultra analyzer equipped with the high-angle collection optics and other modifications of the present invention, as previously described. The lysate supply line was completely disconnected, and the whole blood sample flowed directly into the cuvette assembly 40 without lysate or other diluents. Both analyzers were equipped with Zeonex windows in each cuvette. This dataset was converted into a Matlab cell array file for use with a Matlab script.

[0053] Prediction model: The next step in the calculations was to create predictive models. Three models were developed for the analysis: one for the COOx parameters tHb and COHb, a second for HHb and MetHb, and a third for tBil. The amount of O2Hb was determined by subtracting COHb, HHb, and MetHb from 100%. The X data array was constructed from terms generated from measured absorbance at wavelengths between 462 nm and 650 nm in 1-nm intervals. The tBil model was developed using the same data set as the COOx model, except that samples with MetHb values ​​greater than 20% were excluded from the model. For each model, five Y predicted values ​​(O2Hb, HHb, COHb, MetHb, tBil) were assigned, and tHb was determined by adding the results for O2Hb, HHb, COHb, and MetHb. The number of Y orthogonal values ​​required was determined by manual optimization of the correlation residuals between the mapping function blood predictions and the reference analyzer values.

[0054] Using the initial calibration dataset, the calibration sequence of the machine learning algorithm is A relationship (mapping function) is established between a matrix of sample properties (the Y matrix) and a matrix of measured absorbance at several wavelengths and possibly other measurements based on absorbance versus wavelength (the X matrix). Once this relationship is established, it is used by the analyzer to predict unknown Y values ​​from new measurements of X for the whole blood sample.

[0055] Table 1 summarizes the settings and inputs used for the optimized models. The X data consisted of absorbance and other terms based on absorbance versus wavelength. In the process of optimizing the models, the absorbance derivative versus wavelength was added. Models for analytes more sensitive to nonlinear scattering effects were built with terms for the square root of absorbance and its derivative. Models for analytes more affected by scattering had a correction term proportional to the fourth power of wavelength. The X vector rows have one value for each wavelength for each of the three absorbance-based terms f, g, and h shown in the table for each model.

[0056] [Table 1]

[0057] The calibration set Y matrix is ​​constructed from the known values ​​of the calibration sample set of n lysed blood samples as follows:

[0058]

number

[0059] where tHb is the total hemoglobin value of the lysed blood sample, COHb is the carboxyhemoglobin value of the lysed blood sample; HHb is the deoxygenated hemoglobin value of the lysed blood sample; MetHb is the methemoglobin level in the lysed blood sample; tBil is the total bilirubin value of the lysed blood sample.

[0060] The X matrix is ​​constructed as follows:

[0061]

number

[0062] where f, g, and h are the absorbance base functions listed in Table 1 versus wavelength, respectively. The matrix X includes contributions from absorbance at various wavelengths. The scope of the invention includes optionally adding other measurements to the calculation to reduce interference effects.

[0063] Once these matrices are formed, they are used as a calibration set and the mapping function is calculated according to the specific procedure of the selected machine learning algorithm. As previously mentioned, conventional partial least squares, linear regression, linear algebra, neural networks, multivariate adaptive regression splines, projection onto latent structure, kernel-based orthogonal projection onto latent structure, or other machine learning mathematics is used with the results obtained from the calibration set of data to determine the empirical relationship (or mapping function) between absorbance values ​​and hemoglobin parameters. Typically, a mathematical package is used to generate the results, and the package generally has the option to select one of the machine learning mathematics known to those skilled in the art. Various mathematical packages exist and include, but are not limited to, Matlab by MatWorks, Natick, MA, in combination with Orange data mining software from Orange Bioinformatics, available via the Internet at orange.biolab.si; "R" by the R Project for Statistical Computing, available via the Internet at www.r-project.org; and Python from the Python Software Foundation, available via the Internet at www.python.org, to name a few.

[0064] It is shown that the method of Kernel-Based Orthogonal Projection to Latent Structure (KOPLS) can be used as one type of machine learning algorithm to generate mapping functions. The explanation and description of KOPLS is best exemplified by the following reference: Johan Trygg and Svante Wold, "Orthogonal projections to latent structures (O-PLS)," J. Chemometrics, vol. 1, pp. 111-114, 2014. 2002;16:119-128;Mattias Rantalainen et al. "Kernel-based orthogonal projections to latent structures (K-OPLS)" J.Chemometrics 2007;21:376-385; and Max Bylesjo et al. "K-OPLS package: Kernel-based orthogonal projections to latent structures for prediction and interpretation in feature space," BMC Bioinformatics 2008, 9:106, the contents of which are incorporated herein by reference. Kernel-based mathematics is useful when dealing with nonlinear behavior in a system by using kernel functions to map original data into a high-dimensional space. While any of the machine learning mathematics discussed above can be used to enable one skilled in the art to implement the present invention, KOPLS has additional advantages over other calculations, such as traditional partial least squares. This is because KOPLS This is because the method not only establishes a relationship between the quantified variations to be determined and the analyte value, but also removes variations in the original data that are not quantified but are always present. These non-quantified variations may be due to analyzer and / or blood effects, such as scattering losses and other interference phenomena that are not explicitly measured. By extracting these non-quantified variations from the data, the method leaves information in the data that can be used to predict measurements.

[0065] Using an initial training data set, the KOPLS model establishes a relationship (mapping function) between a matrix of known sample properties (the Y matrix) and a matrix of measured absorbance at several wavelengths and possibly other measurements based on absorbance versus wavelength (the X matrix), processed through a kernel function specified by the KOPLS method. Once the KOPLS coefficients for this relationship are established, they are used by the analyzer in conjunction with the kernel function to predict unknown hemoglobin parameter values ​​from new measurements of absorbance for the sample.

[0066] The kernel function used in this example is described in the above-cited reference by Mattias Rantalainen et al. and has the following formula:

[0067]

number

[0068] where the matrix of measurements X is placed into the kernel function and undergoes further processing as specified in the above-cited KOPLS reference (incorporated by reference) to produce the KOPLS training coefficients.

[0069] Once the set of training coefficients or mapping function is established, it is used to predict hemoglobin and / or total bilirubin parameter values ​​of a blood sample from further measurements. A single-row X matrix is ​​created from the new measurements, and then values ​​from this single-row X matrix are passed through the kernel and mapping function to generate hemoglobin and / or total bilirubin parameter values ​​according to the procedure required for the mapping function used in accordance with the KOPLS procedure, as described in detail in the KOPLS references disclosed above.

[0070] The data collected from the blood samples described above was subjected to a cross-validation process via the KOPLS method. Cross-validation is a process for testing a method using a data set. Some rows of data are set aside, and the remainder are used to create a mapping function. The set aside values ​​are then used as "new" measurements, and their Y matrix values ​​are calculated. This process is repeated by setting aside other measurements and calculating another mapping function. By plotting known values ​​of the blood data against the calculated values, the validity of the method can be confirmed by examining the plot.

[0071] Referring now to Figures 18-23, graphical plots of correlation results comparing various hemoglobin parameters of lysed blood with whole blood using the KOPLS method are shown. Blood samples were manipulated to cover a wide range of analyte values. An n-fold cross-validation technique using 60 folds was used to test the data. In this technique, the data set is divided into n=60 distinct sets, a model is created from n-1 of the sets, and the remaining sets are predicted using the model. The process is repeated 60 times for each group, so that all data points are predicted without being included in the model. The predictions are made using a model built from most of the data.

[0072] Figure 19 shows correlation results for tHb using the K-OPLS method. The horizontal axis has units representing total bilirubin in grams per deciliter of lysed blood. The vertical axis has units representing total bilirubin in grams per deciliter of whole blood. As can be seen from the plot, the method for determining tHb ​​in whole blood samples has a correlation of greater than 99%.

[0073] 20 shows correlation results for O2Hb using the K-OPLS method. The horizontal axis has units representing percent oxygenated hemoglobin of lysed blood. The vertical axis has units representing percent oxygenated hemoglobin of whole blood. As can be seen from the plot, the method for determining O2Hb in whole blood samples has a correlation of greater than 99%.

[0074] 21 shows correlation results for carboxyhemoglobin using the K-OPLS method. The horizontal axis has units representing percent carboxyhemoglobin of lysed blood. The vertical axis has units representing percent carboxyhemoglobin of whole blood. As can be seen from the plot, the method for determining COHb in whole blood samples has a correlation of greater than 99%.

[0075] 22 shows correlation results for deoxyhemoglobin using the K-OPLS method. The horizontal axis has units representing percent deoxyhemoglobin of lysed blood. The vertical axis has units representing percent deoxyhemoglobin of whole blood. As can be seen from the plot, the method for determining HHb in whole blood samples has a correlation of greater than 99%.

[0076] 23 shows correlation results for methemoglobin using the K-OPLS method. The horizontal axis has units representing percent methemoglobin in lysed blood. The vertical axis has units representing percent methemoglobin in whole blood. As can be seen from the plot, the method for determining MetHb in whole blood samples has a correlation of greater than 99%.

[0077] Figure 24 shows correlation results for tBil using the K-OPLS method. The horizontal axis has units representing total bilirubin in milligrams per deciliter of lysed blood. The vertical axis has units representing total bilirubin in milligrams per deciliter of whole blood. As can be seen from the plot, the method for determining tBil of whole blood samples has a correlation of greater than 99%.

[0078] A method for performing whole blood measurements using the COOx analyzer system 10 of the present invention is now described. An absorbance scan is measured by first recording a transmitted light intensity scan with the cuvette module 43 filled with a transparent fluid, such as water or analyzer flush solution, known as a "blank" scan. Then, a transmitted light intensity scan with the cuvette module 43 filled with a whole blood sample is recorded. After correcting for spectrometer dark response and detector linearity, the spectral absorbance is the negative logarithm to the base 10 of the ratio of the whole blood scan to the transmitted fluid scan calculated at each wavelength within the measurement range.

[0079] More specifically, diagrams of the components of the COOx analyzer subsystem are shown in Figures 1-18. This embodiment of the subsystem measures the absorbance of a liquid introduced into a cuvette module 43. The light used to perform the absorbance measurement originates from an LED light source 28, is collected and transmitted by a collimating lens 30, and passes through a first diffuser plate 32, a circular polarizer 34, a converging lens 36, and an optional protective window 38 before reaching the cuvette module 43. Knowledge of the cuvette path length is important for the absorbance measurement. The cuvette path length is pre-measured for each individual cuvette module 43 and programmed into an electronic chip 48c on the cuvette module 43. The path length information is Whenever needed, it is read / retrieved by the analyzer's data processor module 130 .

[0080] After passing through the cuvette module 43, the light is collected and collimated by a lens 66 and sent through a second diffuser 68 and a beam splitter 69. The purpose of the beam splitter 69 is to allow light from a calibration light source 72 (e.g., a krypton gas discharge lamp) to enter the light path 21, where it is collected by a lens 74. The calibration light source 72 provides light of a few known wavelengths, which is used to periodically calibrate the wavelength scale of the spectrometer module 100. After passing through the beam splitter 69, the light is focused by a lens 82 onto an optical fiber 92. The optical fiber 92 directs the light to the input slit 114 of the spectrometer module 100. The light passes through an achromatic lens 124 and travels through an optically dispersive element 130 having a reflective back portion 132. The light is wavelength dispersed by passing through a light dispersing element 130, such as a prism 130, and then makes a return pass through a lens 124, which refocuses the light onto the pixels of the optical array detector 116. The optical array detector 116 converts the optical energy into electrical signals indicative of the spectral intensity of the light. The electrical signals are sent to a data processor module 150 for further processing and display of the final results to a user. The optical receiving and converting assembly 110 is a single substrate that holds the input slit 114 and the optical array detector 116 in close proximity as an integral unit.

[0081] The input slit 114 is applied directly on the same circuit board substrate 112 as and in close proximity to the optical array detector 116. Other prior art spectrometers place these components on separate planes, with the components having separate mounting structures that require independent adjustment and alignment. The mounting scheme of the present invention has several advantages that reduce the cost and size of the spectrometer module 100. These advantages include: 1) the cost of a separate mounting structure is avoided; 2) the input slit 114 can be laser etched in a precise location relative to the optical array detector 116, making alignment less labor-intensive; 3) inexpensive spherical surface optics can be used in the optical system because the image of the slit on the detector is only slightly off-axis from the central axis of the optical system, minimizing aberrations; and 4) a single alignment procedure for the integrated slit and detector assembly replaces alignment procedures for two separate assemblies.

[0082] It is important to note that the first and second diffusers 32 and 68 are positioned in front of and behind the cuvette module 43, respectively. Measuring the absorbance of diffuse samples presents unique challenges. The diffuse transmission of the sample perturbs the initial spatial light distribution of the measurement system, which is caused by the inherent non-uniformity of the light source. Therefore, the spatial light distribution of a "blank" scan may be quite different from that of a whole-blood sample scan. Because optical detectors have spatially varying responses, even if the overall intensity remains unchanged, the response may fluctuate with changes in the spatial distribution of the incident light. An absorbance scan based on the ratio of the sample scan to the blank scan will have a significant absorbance component due to this effect in addition to the absorbance due to the sample alone. This results in significant measurement error in sample absorbance, which is unacceptable for co-oximetry.

[0083] The advantage of placing the cuvette module 43 between the first and second diffusers 32 and 68 is that the spatial light distribution will look the same for the blank and sample scans, eliminating this error effect. The diffusers 32, 68 are specifically selected to diffuse the incoming light rays so that they fall into the entire acceptance cone of the optical system, but not so much so that as much light throughput as possible is preserved while perturbing light rays that traverse the field completely.

[0084] While preferred embodiments of the present invention have been described herein, the foregoing description is by way of example only. Further modifications of the invention disclosed herein will occur to those skilled in the art, and all such modifications are deemed to be within the scope of the invention as defined by the scope of the appended patent applications.

Claims

1. A system (10) for measuring whole blood hemoglobin parameters, comprising: An optical sample module (20), comprising: a light emitting module (22) having an LED light source (28) capable of emitting light, the light being directed thereby defining an optical path (21); a replaceable cuvette assembly (40) adjacent to the light-emitting module (22), the replaceable cuvette assembly (40) being adapted to receive a whole blood sample, the replaceable cuvette assembly (40) having a sample-receiving chamber (54) having a first cuvette window (49) and a second cuvette window (52) aligned with the first cuvette window (49), the sample-receiving chamber (54) being disposed in the light path (21) to receive light from the LED light source (28), the sample-receiving chamber (54) having a defined optical path length (43 a) between the first cuvette window (49) and the second cuvette window (52), and an electronic chip (48 c) capable of storing a path length value of the sample-receiving chamber (54); and an optical sample module (20) including a calibration light module (60) having a calibration light source (72) having light of one or more known wavelengths, the calibration light module (60) being capable of emitting calibration light into the optical path (21); an optical fiber (92) having a light-receiving end (92a) and a light-emitting end (92b), the light-receiving end (92a) optically connected to the optical sample module (20) to receive light emitted along the optical path (21) and transmit the light to the light-emitting end (92b); a spectrometer module (100) capable of receiving light from the light-emitting end (92b) of the optical fiber (92), separating the light into a plurality of light beams each having a different wavelength, and converting the plurality of light beams into electrical signals; a processor module (150) capable of obtaining (1) a pathlength value of the sample receiving chamber (54) from the electronic chip (48c), receiving and processing (2) the electrical signals from the spectrometer module (100) generated for the whole blood sample, and converting the electrical signals, depending on the pathlength value of the sample receiving chamber (54), into output signals usable for displaying and reporting hemoglobin and / or total bilirubin parameter values ​​for the whole blood sample.

2. 2. The system (10) of claim 1, wherein the light-emitting module (22) includes a plurality of optical components (B) disposed in an optical path (21) between the LED light source (28) and the cuvette assembly (40), the plurality of optical components (B) including at least a first optical diffuser (32), and one or more of a collimating lens (30), a circular polarizer (34), and a converging lens (36).

3. 2. The system of claim 1, wherein the calibration light module includes a second optical diffuser disposed in the optical path downstream from the cuvette assembly but upstream from a beam splitter.

4. An absorbance measurement system (10) for whole blood, comprising: An optical sample module (20), comprising: a light emitting module (22) having an LED light source (28) capable of emitting light, the light being directed thereby defining an optical path (21); a replaceable cuvette module (43) adjacent to the light-emitting module (22), the sample receiving module being adapted to receive a whole blood sample and having a first cuvette window (49) and a second cuvette window (52) aligned with the first cuvette window (49); a replaceable cuvette module (43) having a sample-receiving chamber (54) disposed in the light path (21) for receiving light from the LED light source (28); a first optical diffuser (32) positioned in the optical path (21) between the LED light source (28) and a replaceable cuvette module (43); and a replaceable cuvette module (43) having a second optical diffuser (68) positioned in the optical path (21) after the replaceable cuvette module (43); an optical fiber (92) having a light-receiving end (92a) and a light-emitting end (92b), the light-receiving end (92a) optically connected to the optical sample module (20) to receive light emitted along the optical path (21) and transmit the light to the light-emitting end (92b); a spectrometer module (100) capable of receiving light from the light-emitting end (92b) of the optical fiber (92), separating the light into a plurality of light beams each having a different wavelength, and converting the plurality of light beams into electrical signals; a processor module (150) capable of obtaining (1) a pathlength value of the sample receiving chamber (54) from the electronic chip (48c), receiving and processing (2) the electrical signals from the spectrometer module (100) generated for the whole blood sample, and using the pathlength value of the sample receiving chamber (54) to convert the electrical signals into output signals usable for displaying and reporting hemoglobin and / or total bilirubin parameter values ​​for the whole blood sample.

5. 5. The absorbance measurement system (10) of claim 4, wherein the light-emitting module (22) includes a plurality of optical components (B) arranged in an optical path (21) between the LED light source (28) and the replaceable cuvette module (43), the plurality of optical components (B) including at least a first optical diffuser (32), and one or more of a collimating lens (30), a circular polarizer (34), and a converging lens (36).

6. The spectrometer module (100) comprises: an input slit (114) positioned in the optical path (21) to receive light emitted from the light-emitting end (92b) of the optical fiber (92) and transmit the light through the input slit (114); an optical component group (120) having an achromatic lens (124) and an optical dispersive element (130), the optical dispersive element (130) disposed in the optical path (21), the optical dispersive element (130) capable of receiving light transmitted through the input slit (114) and the achromatic lens (124), separating the light into multiple light beams each having a different wavelength, and redirecting the multiple light beams back through the achromatic lens (124) toward but offset from the input slit (114); an optical array detector capable of receiving the plurality of light beams and converting the plurality of light beams into electrical signals.

7. An optical spectrometer (100) for use in a COOx analyzer, comprising: a spectrometer housing (102) having a base plate (104a) and an optical fiber receiving end (108); a receiving input slit (114) positioned adjacent the optical fiber receiving end (108) of the spectrometer housing (102); A spectrometer housing (102) is positioned away from the optical fiber receiving end (108) and within the optical path (21) along which light travels from the receiving input slit (114). an optically dispersive element (130) mounted on the base plate (104a) and capable of receiving light transmitted through the input slit (114), separating the light into multiple light beams each having a different wavelength, and redirecting the multiple light beams; an optical array detector (116) capable of receiving the plurality of light beams and converting the plurality of light beams into electrical signals; an achromatic lens (121) having a lens mount (122) and an achromatic lens (124) mounted in the lens mount (122), the achromatic lens (124) positioned in the optical path (21) to direct light from the input slit (114) to the light dispersive element (130), receive a plurality of light beams reflected from the light dispersive element (130), and direct the plurality of light beams onto the optical array detector (116); and a thermal compensation means for maintaining the positions of the plurality of light beams on the optical array detector, the thermal compensation means comprising one or more of a thermal insulator disposed around the spectrometer housing, a temperature controller assembly, and a thermal compensation lens mount.

8. 8. The optical spectrometer of claim 7, wherein the thermal compensation lens mount has a fixed mount end and a non-fixed mount end that allows for thermal expansion and contraction of the thermal compensation lens mount, the fixed mount end being fixedly attached to the base plate, and the lens mount having a coefficient of expansion greater than the coefficient of expansion of the base plate.

9. 8. The optical spectrometer (100) of claim 7, wherein the thermal compensation lens mount (122) moves linearly and laterally relative to the optical path (21) of light from the optical input slit (114) based on an expansion coefficient of the lens mount (122) to maintain the position of dispersed light from the optical dispersion element (130) on the optical array detector (116).

10. A compact spectrometer (100) for measuring hemoglobin parameters in whole blood, comprising: an enclosed spectrometer housing (102) having an optical fiber receiving end (108) with an optical input port (109); an optical input slit (114) disposed on a circuit board substrate (112) disposed within the enclosed spectrometer housing (102), the optical input slit (114) being aligned with and adjacent to the optical input port (109); an optical array detector (116) disposed on the circuit board substrate (112) adjacent to the optical input slit (114); a group of optical components (120) consisting of a light-dispersing element (130) disposed downstream from the light input slit (114) and a spherical achromatic lens (120) disposed between the light input slit (114) and the light-dispersing element (130), the light-dispersing element (130) having a reflective surface (132), and the achromatic lens (120) transmitting light from the light input slit (114) to the light-dispersing element (130) and transmitting dispersed light reflected from the light-dispersing element (130) to the light array detector (116).

11. 11. The compact spectrometer of claim 10, further comprising a thermal compensation means (160) for maintaining the positions of the plurality of light beams on the optical array detector (116), the thermal compensation means (160) comprising one or more of a thermal insulator disposed around the spectrometer housing (102), a temperature controller assembly (170), and / or a thermal compensation lens mount (122). １００）。

12. 12. The compact spectrometer of claim 11, wherein the thermal compensation lens mount has a fixed mount end and a non-fixed mount end that allows for thermal expansion and contraction of the thermal compensation lens mount, the fixed mount end being fixedly attached to the base plate, and the lens mount having a coefficient of expansion greater than the coefficient of expansion of the base plate.

13. 1. A method for measuring whole blood hemoglobin parameters despite strong optical scattering caused by whole blood, comprising: providing an LED light source (28) having a spectral range of about 422 nm to about 695 nm; directing light having said spectral range along an optical path (21); providing a cuvette module (43) having a sample receiving chamber (54) with a first cuvette window (49) and a second cuvette window (52) disposed in the optical path (21), the second cuvette window (52) of the cuvette module (43) transmitting light to the sample receiving chamber (54) and the first cuvette window (49), the sample receiving chamber (54) containing a sample of whole blood; providing a pair of first and second optical diffusers (32, 68) disposed in the optical path (21), wherein the first cuvette window (49) and the second cuvette window (52) of the sample receiving chamber (54) of the cuvette module (43) are disposed between the pair of first and second optical diffusers (32, 68); directing the light from the cuvette module (43) into a spectrometer (100) having a light dispersing element (130) that separates the light into multiple light beams, each having a different wavelength, and converts the multiple light beams into electrical signals; and processing the electrical signal into an output signal usable for displaying and reporting a hemoglobin parameter value and / or a total bilirubin parameter value of the sample of whole blood.

14. The processing step includes:

14. The method of claim 13, further comprising processing the electrical signals into spectral absorptances and then mapping the spectral absorptances to hemoglobin and / or total bilirubin parameter values ​​using a computational mapping function.

15. The method of claim 14 , wherein the processing step includes using a kernel-based orthogonal projection to latent structure mapping function as the computational mapping function.

16. 1. A method for measuring a hemoglobin parameter in a whole blood sample using absorbance, comprising: measuring and recording transmitted light intensity scans across multiple wavelengths within a measurement range by transmitting light through a cuvette module (43), the cuvette module (43) having an optical path (21) with a known optical path length through the cuvette module (43) and filled with a transparent fluid; and measuring and recording a transmitted light intensity scan across the plurality of wavelengths in the measurement range by transmitting light a second time through the cuvette module (43), the cuvette module (43) having the light path (21) with the known optical path length through the cuvette module (43) and filled with a whole blood sample, wherein the steps of measuring and recording the transmitted light intensity scan across the plurality of wavelengths in the measurement range are performed by transmitting the transmitted light a second time through the cuvette module (43), the cuvette module (43) having the light path (21) with the known optical path length through the cuvette module (43). diffusing and circularly polarizing the transmitted light before transmitting it to a cuvette module (43); and subsequently diffusing the transmitted light exiting the cuvette module (43) before determining the spectral absorptance, determining a spectral absorptance at each of the plurality of wavelengths in the measurement range based on a ratio of the transmitted light intensity scan of the whole blood sample to the transmitted light intensity scan of the transparent fluid using a spectrometer (100); and and correlating the absorbance at each wavelength of the plurality of wavelengths in the measurement range with hemoglobin and / or bilirubin parameter values ​​of the blood sample using a computational mapping function.

17. 17. The method of claim 16, further comprising controlling temperature-induced drift of the transmitted light within the spectrometer (100) by one or more of insulating the spectrometer (100), controlled heating of the spectrometer (100), or incorporating a temperature-compensating lens mount (122) within the spectrometer (100).

18. 17. The method of claim 16, wherein incorporating a temperature-compensating lens mount (122) includes linearly and laterally moving the optical path (21) of the light from the light input slit (114) based on an expansion coefficient of the lens mount (122), thereby maintaining or moving the position of the transmitted light from the light dispersive element (130) on the optical array detector (116).

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