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13 pages, 4616 KiB  
Article
Non-Invasive Glucose Measurement Technique Based on Time-of-Flight
by Konstantinos Asimakopoulos and Evangelos Hristoforou
Appl. Sci. 2024, 14(24), 11602; https://doi.org/10.3390/app142411602 - 12 Dec 2024
Viewed by 917
Abstract
A novel, low-cost sensor for the measurement of physiological levels of glucose in samples has been developed using Time-of-Flight (ToF) technology. The sensor, built with off-the-shelf components, leverages the scattering of light by glucose molecules to determine concentration changes by measuring the phase [...] Read more.
A novel, low-cost sensor for the measurement of physiological levels of glucose in samples has been developed using Time-of-Flight (ToF) technology. The sensor, built with off-the-shelf components, leverages the scattering of light by glucose molecules to determine concentration changes by measuring the phase shift of light signals propagating through the sample. Experimental validation demonstrates the feasibility of this approach, correlating glucose concentration with optical scattering effects at a wavelength of 850 nm. This technique offers the potential for a cost-effective, non-invasive alternative to traditional methods of glucose monitoring, with applications in continuous glucose monitoring systems to improve diabetes management by reducing patient discomfort and enhancing monitoring accuracy. Further optimization of the sensor parameters and expanded testing are still needed. Full article
(This article belongs to the Section Applied Physics General)
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<p>Simplified schematic of a heterodyne ToF sensor.</p>
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<p>Representation of scattering of light from glucose molecules. Changing concentration affects sample scattering.</p>
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<p>The experimental setup developed.</p>
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<p>OTP8320 schematic.</p>
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<p>Phase data from sensor, rendered as a grayscale image.</p>
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<p>High-level diagram of the developed circuit around to operate the sensor.</p>
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<p>A 3D rendering of the designed sample holder.</p>
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<p>Averaged sensor output (phase of received signal) versus sensor pixel for different glucose concentrations in the sample.</p>
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<p>Averaged sensor output (amplitude of received signal) versus sensor pixel for different glucose concentrations in the sample.</p>
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<p>Averaged sensor output (phase channel) for different glucose concentrations.</p>
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<p>Averaged sensor output (amplitude channel) for different glucose concentrations.</p>
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<p>Averaged sensor output (phase channel scaled by amplitude channel) for different glucose concentrations.</p>
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12 pages, 7395 KiB  
Article
Multi-Cavity Nanorefractive Index Sensor Based on MIM Waveguide
by Weijie Yang, Shubin Yan, Ziheng Xu, Changxin Chen, Jin Wang, Xiaoran Yan, Shuwen Chang, Chong Wang and Taiquan Wu
Nanomaterials 2024, 14(21), 1719; https://doi.org/10.3390/nano14211719 - 28 Oct 2024
Viewed by 685
Abstract
Within this manuscript, we provide a novel Fano resonance-driven micro-nanosensor. Its primary structural components are a metal-insulator-metal (MIM) waveguide, a shield with three disks, and a T-shaped cavity (STDTC). The finite element approach was used to study the gadget in theory. It is [...] Read more.
Within this manuscript, we provide a novel Fano resonance-driven micro-nanosensor. Its primary structural components are a metal-insulator-metal (MIM) waveguide, a shield with three disks, and a T-shaped cavity (STDTC). The finite element approach was used to study the gadget in theory. It is found that the adjustment of the structure and the change of the dimensions are closely related to the sensitivity (S) and the quality factor (FOM). Different model structural parameters affect the Fano resonance, which in turn changes the transmission characteristics of the resonator. Through in-depth experimental analysis and selection of appropriate parameters, the sensor sensitivity finally reaches 3020 nm/RIU and the quality factor reaches 51.89. Furthermore, the installation of this microrefractive index sensor allows for the quick and sensitive measurement of glucose levels. It is a positive contribution to the field of optical devices and micro-nano sensors and meets the demand for efficient detection when applied in medical and environmental scenarios. Full article
(This article belongs to the Section Nanoelectronics, Nanosensors and Devices)
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<p>Two-dimensional schematic diagram of a shield-shaped structure with three discs and a T-shaped cavity.</p>
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<p>(<b>a</b>) Transmission spectra of different structures; (<b>b</b>) Shield structure with T-cavity.</p>
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<p>Magnetic field strength distribution (<b>a</b>) Single shield cavity; (<b>b</b>) Shield with T-cavity; (<b>c</b>) STDTC.</p>
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<p>(<b>a</b>) Transmission curves with varying R<sub>1</sub> lengths; (<b>b</b>) Sensitivity fit lines with varying R<sub>1</sub> lengths; (<b>c</b>) Comparing FWHM with FOM for various R<sub>1</sub> lengths.</p>
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<p>(<b>a</b>) Transmission curves with varying L lengths; (<b>b</b>) Sensitivity fit lines with varying L lengths; (<b>c</b>) Comparing FWHM with FOM for various L lengths.</p>
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<p>(<b>a</b>) Transmission curves with varying r<sub>1</sub> lengths; (<b>b</b>) Sensitivity fit lines with varying r<sub>1</sub> lengths.</p>
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<p>(<b>a</b>) Transmission curves with varying g lengths; (<b>b</b>) Comparing FWHM with FOM for various g lengths.</p>
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<p>(<b>a</b>) Transmission spectra for various indexes of refraction; (<b>b</b>) Refractive index sensitivity fit lines.</p>
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<p>(<b>a</b>) Transmission spectrum of glucose concentration; (<b>b</b>) Fitted line of glucose concentration.</p>
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59 pages, 4638 KiB  
Review
Cutting-Edge Hydrogel Technologies in Tissue Engineering and Biosensing: An Updated Review
by Nargish Parvin, Vineet Kumar, Sang Woo Joo and Tapas Kumar Mandal
Materials 2024, 17(19), 4792; https://doi.org/10.3390/ma17194792 - 29 Sep 2024
Viewed by 1757
Abstract
Hydrogels, known for their unique ability to retain large amounts of water, have emerged as pivotal materials in both tissue engineering and biosensing applications. This review provides an updated and comprehensive examination of cutting-edge hydrogel technologies and their multifaceted roles in these fields. [...] Read more.
Hydrogels, known for their unique ability to retain large amounts of water, have emerged as pivotal materials in both tissue engineering and biosensing applications. This review provides an updated and comprehensive examination of cutting-edge hydrogel technologies and their multifaceted roles in these fields. Initially, the chemical composition and intrinsic properties of both natural and synthetic hydrogels are discussed, highlighting their biocompatibility and biodegradability. The manuscript then probes into innovative scaffold designs and fabrication techniques such as 3D printing, electrospinning, and self-assembly methods, emphasizing their applications in regenerating bone, cartilage, skin, and neural tissues. In the realm of biosensing, hydrogels’ responsive nature is explored through their integration into optical, electrochemical, and piezoelectric sensors. These sensors are instrumental in medical diagnostics for glucose monitoring, pathogen detection, and biomarker identification, as well as in environmental and industrial applications like pollution and food quality monitoring. Furthermore, the review explores cross-disciplinary innovations, including the use of hydrogels in wearable devices, and hybrid systems, and their potential in personalized medicine. By addressing current challenges and future directions, this review aims to underscore the transformative impact of hydrogel technologies in advancing healthcare and industrial practices, thereby providing a vital resource for researchers and practitioners in the field. Full article
(This article belongs to the Special Issue Advanced Composite Biomaterials for Tissue Regeneration)
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<p>Illustration depicting different designs of hydrogel biosensors and tissue engineering applications, along with their respective functional roles.</p>
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<p>(<b>a</b>,<b>b</b>) Overview of the physical and mechanical characteristics of (FL) hydrogels. Reprinted with permission from [<a href="#B35-materials-17-04792" class="html-bibr">35</a>]. <a href="https://doi.org/10.1038/ncomms3974" target="_blank">https://doi.org/10.1038/ncomms3974</a>. (<b>c</b>) Diagram illustrating the different types of physical and chemical cross-linking methods used in hydrogels (adapted from [<a href="#B36-materials-17-04792" class="html-bibr">36</a>], CC-BY license) <a href="https://doi.org/10.3390/gels7040255" target="_blank">https://doi.org/10.3390/gels7040255</a>.</p>
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<p>The fabrication process of Gel/PP-TA-Ag hydrogel, designed for wound healing with hemostatic and antibacterial properties. Reprinted with permission from [<a href="#B44-materials-17-04792" class="html-bibr">44</a>]. <a href="https://doi.org/10.1016/j.nantod.2021.101165" target="_blank">https://doi.org/10.1016/j.nantod.2021.101165</a>.</p>
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<p>Diagram showing the steps involved in creating macroporous hydrogel scaffolds using sacrificial porogens. (A) Degradable particles are gathered and dispersed in a cross-linkable prepolymer solution. (B) The mixture, containing the porogens (in this case, microspheres), is poured into a mold before solidification. (C) Gelation occurs as the polymerization process encapsulates the leachable porogens. (D) Controlled degradation of the porogens results in the formation of a three-dimensional porous network with interconnected macropores and channels. Reprinted with permission from [<a href="#B47-materials-17-04792" class="html-bibr">47</a>]. <a href="https://doi.org/10.5051/jpis.2013.43.6.251" target="_blank">https://doi.org/10.5051/jpis.2013.43.6.251</a>.</p>
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<p>(<b>1</b>) Overview of the submerged bioprinting technique. (<b>A</b>) Cell-laden hydrogel droplets are sequentially deposited layer by layer following a predefined model to create a three-dimensional tissue structure. This process takes place within a high-density fluorocarbon liquid that provides buoyant support. (<b>B</b>) Hydrogel droplets can be added either vertically or horizontally to an existing structure. The fluorocarbon support allows for the creation of branching hydrogel structures or cantilever-like formations without needing a solid foundation. (<b>2</b>) Dispensing accuracy. (<b>A</b>) To evaluate the precision of the dispensing process, the offset distance (Δ) from the needle’s center to the printed drop was measured, along with the drop diameter. The dispensing deviation was calculated using Equation (1). Printing accuracy was compared between printing in air (<b>B</b>) and printing in FC-43 (<b>C</b>), with submerged printing showing improved accuracy. The findings are displayed in boxplot diagrams, <span class="html-italic">n</span> = 20. Reprinted with permission from [<a href="#B57-materials-17-04792" class="html-bibr">57</a>]. <a href="https://doi.org/10.1089/biores.2013.0031" target="_blank">https://doi.org/10.1089/biores.2013.0031</a>.</p>
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<p>Techniques for designing material, cell, and tissue architectures aimed at tissue engineering applications. Key scaffold requirements for bone tissue engineering are also highlighted. Reprinted with permission from [<a href="#B75-materials-17-04792" class="html-bibr">75</a>]. <a href="https://doi.org/10.3390/ma14226899" target="_blank">https://doi.org/10.3390/ma14226899</a>.</p>
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<p>Hierarchical organization of hydrogel structures. The structure of hydrogels can be systematically organized across multiple scales and levels of information content. Starting with material selection, hydrogels can be structured hierarchically from the first to the third order by increasing scale, and from the fourth to the sixth order by adding complexity in information content. The highest level (sixth order) represents a precisely controlled environment, both spatially and temporally, that can mimic the patterning cues observed during neural development. Reprinted with permission from [<a href="#B87-materials-17-04792" class="html-bibr">87</a>]. <a href="https://doi.org/10.1016/j.biotechadv.2019.03.009" target="_blank">https://doi.org/10.1016/j.biotechadv.2019.03.009</a>.</p>
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<p>(<b>Top</b>) Diagram showing the preparation of hydrogel composites through the in situ synthesis of PBNPs within the CMC matrix in the presence of enzymes (Gox or ADH). The successful formation of the hydrogel composite is demonstrated by a tube inversion test (central image). Modifying flexible screen-printed electrodes (SPE) with the hydrogel composite produces a porous film on the electrode surface, confirmed by SEM and optical microscopy images (inset) of the PBNPs-CMC hydrogel. (<b>Bottom</b>) The operational principle of the ethanol biosensor relies on the electrocatalytic oxidation of NADH, generated by ADH-mediated ethanol oxidation on the PBNPs-ADH-CMC modified SPE, while the glucose biosensor is based on the electrocatalytic reduction of H2O2, produced by Gox-mediated glucose oxidation on the PBNPs-Gox-CMC modified SPE. (<b>A</b>) Calibration curve for increasing hydrogen peroxide concentrations on PBNPs-CMC modified SPE in 0.05 M phosphate buffer (pH 7.4) with 150 mM KCl. Inset: chronoamperometric curves at −0.1 V (vs. Ag/AgCl). (<b>B</b>) Calibration curve for increasing glucose concentrations on PBNPs-Gox-CMC modified SPE in human serum samples. Inset: chronoamperometric curves at −0.1 V (vs. Ag/AgCl). Reprinted with permission from [<a href="#B108-materials-17-04792" class="html-bibr">108</a>]. <a href="https://doi.org/10.1016/j.snb.2022.132985" target="_blank">https://doi.org/10.1016/j.snb.2022.132985</a>.</p>
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<p>Design and development of advanced composite hydrogels tailored for wearable health monitoring applications. Reprinted with permission from [<a href="#B131-materials-17-04792" class="html-bibr">131</a>]. <a href="https://doi.org/10.1007/s40820-023-01079-5" target="_blank">https://doi.org/10.1007/s40820-023-01079-5</a>.</p>
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<p>Investigating the function of magnetic nanoparticle-infused hydrogels in diagnostic applications. Reprinted with permission from [<a href="#B96-materials-17-04792" class="html-bibr">96</a>]. <a href="https://doi.org/10.1039/D3RA07391B" target="_blank">https://doi.org/10.1039/D3RA07391B</a>.</p>
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<p>Diagram depicting a flexible gas sensor device and its use in detecting bad breath, monitoring meat freshness, and alerting for hydrogen sulfide leaks. Reprinted with permission from [<a href="#B153-materials-17-04792" class="html-bibr">153</a>]. <a href="https://doi.org/10.1038/s41467-023-40953-z" target="_blank">https://doi.org/10.1038/s41467-023-40953-z</a>.</p>
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<p>(<b>a</b>) Tensile and compressive properties of AAV-2 hydrogel before and after freezing, (<b>b</b>) photographs showing AAV-2 hydrogel bent into ‘2’ and ‘U’ shapes post-freezing, (<b>c</b>) electrical circuit using AAV-2 hydrogel as a conductor at 25 °C and −20 °C, (<b>d</b>) current response during the healing cycle of a cut in frozen AAV-2 hydrogel (repeated 5 times), (<b>e</b>) real-time finger bending response monitored by AAV hydrogel at −20 °C. Reprinted with permission from [<a href="#B172-materials-17-04792" class="html-bibr">172</a>].</p>
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12 pages, 5829 KiB  
Article
High-Sensitivity Janus Sensor Enabled by Multilayered Metastructure Based on the Photonic Spin Hall Effect and Its Potential Applications in Bio-Sensing
by Xiang Li and Haifeng Zhang
Sensors 2024, 24(17), 5796; https://doi.org/10.3390/s24175796 - 6 Sep 2024
Viewed by 736
Abstract
The refractive index (RI) of biological tissues is a fundamental material parameter that characterizes how light interacts with tissues, making accurate measurement of RI crucial for biomedical diagnostics and environmental monitoring. A Janus sensor (JBS) is designed in this paper, and the photonic [...] Read more.
The refractive index (RI) of biological tissues is a fundamental material parameter that characterizes how light interacts with tissues, making accurate measurement of RI crucial for biomedical diagnostics and environmental monitoring. A Janus sensor (JBS) is designed in this paper, and the photonic spin Hall effect (PSHE) is used to detect subtle changes in RI in biological tissues. The asymmetric arrangement of the dielectric layers breaks spatial parity symmetry, resulting in significantly different PSHE displacements during the forward and backward propagation of electromagnetic waves, thereby realizing the Janus effect. The designed JBS can detect the RI range of 1.3~1.55 RIU when electromagnetic waves are incident along the +z-axis, with a sensitivity of 96.29°/refractive index unit (RIU). In the reverse direction, blood glucose concentrations are identified by the JBS, achieving a sensitivity of 18.30°/RIU. Detecting different RI range from forward and backward scales not only overcomes the limitation that single-scale sensors can only detect a single RI range, but also provides new insights and applications for optical biological detection through high-sensitivity, label-free and non-contact detection. Full article
(This article belongs to the Special Issue Feature Papers in Biosensors Section 2024)
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<p>A schematic representation of the JBS is described as (SiO<sub>2</sub>·Analyte·SiO<sub>2</sub>·Plasma)<span class="html-italic"><sup>N</sup></span><sup>1</sup>·SiO<sub>2</sub>, where the number of cycles <span class="html-italic">N</span><sub>1</sub> = 8. The upper right corner is the schematic of the photon spin Hall effect, and the lower right corner is the physical parameters of the dielectric layer.</p>
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<p>The Fresnel reflection coefficient of (<b>a</b>) <span class="html-italic">f</span> = 70 GHz, (<b>b</b>) <span class="html-italic">f</span> = 71 GHz, (<b>c</b>) <span class="html-italic">f</span> = 72 GHz, <span class="html-italic">P</span>-wave and <span class="html-italic">S</span>-wave at different EMWs incident frequencies. (<b>d</b>) Fresnel reflection coefficient ratio of <span class="html-italic">P</span>-waves and <span class="html-italic">S</span>-waves.</p>
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<p>(<b>a</b>) cos (<span class="html-italic">φ</span><sub>p</sub> − <span class="html-italic">φ</span><sub>s</sub>) and (<b>b</b>) δ<sub>V</sub><sup>−</sup>/λ with different <span class="html-italic">θ</span> at 70, 71 and 72 GHz.</p>
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<p>The effect of (<b>a</b>) <span class="html-italic">d</span><sub>Analyte</sub>, (<b>b</b>) <span class="html-italic">d</span><sub>SiO2</sub> and (<b>c</b>) <span class="html-italic">d</span><sub>Plasma</sub> changes on δ<sub>V</sub><sup>−</sup>/λ, and (<b>d</b>) the effect of the number of media cycles <span class="html-italic">N</span><sub>1</sub> on δ<sub>V</sub><sup>−</sup>/λ.</p>
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<p>(<b>a</b>) The enhancement effect of the plasma. (<b>b</b>) The transmission and reflectance of the JBS.</p>
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<p>(<b>a</b>) The δ<sub>V</sub><sup>−</sup>/λ of <span class="html-italic">n</span><sub>Analyte</sub> for forward propagation. (<b>b</b>) Fit of angle of incidence and <span class="html-italic">n</span><sub>Analyte</sub> during forward propagation.</p>
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<p>(<b>a</b>) Backward EMW propagation to detect <span class="html-italic">n</span><sub>Analyte</sub>. (<b>b</b>) Fitting of the incidence angle and the blood glucose concentration.</p>
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12 pages, 1331 KiB  
Article
Multimodal Interference-Based Fiber Optic Sensors for Glucose and Moisture Content Detection in Honey
by Mayeli Anais Pérez-Rosas, Yahir Nicolás García-Guevara, Yadira Aracely Fuentes-Rubio, René Fernando Domínguez-Cruz, Oscar Baldovino-Pantaleón and Gerardo Romero-Galván
Appl. Sci. 2024, 14(17), 7914; https://doi.org/10.3390/app14177914 - 5 Sep 2024
Cited by 1 | Viewed by 926
Abstract
Fiber optic sensors (FOSs) have transformed industrial applications with their high sensitivity and precision, especially in real-time monitoring. This study presents a fiber optic sensor based on multimodal interference (MMI) applied to detect honey adulteration. The sensor is built using a non-core multimode [...] Read more.
Fiber optic sensors (FOSs) have transformed industrial applications with their high sensitivity and precision, especially in real-time monitoring. This study presents a fiber optic sensor based on multimodal interference (MMI) applied to detect honey adulteration. The sensor is built using a non-core multimode fiber (NC-MMF) segment spliced between two standard single-mode fibers (SMFs). We focus on reporting the detection of two main adulterants in honey that modify its refractive index (RI): the presence of glucose and moisture content. Detailed testing was performed with two commercially approved honey brands, named A and B. The sensor successfully detected glucose concentrations from 1% to 5% and moisture content from 0% to 20% for both brands. For glucose detection, we obtained sensitivity values −0.55457 nm/% for brand A and −2.61257 nm/% for brand B. In terms of moisture content in honey, we observed a sensitivity around −0.3154 nm/% and −0.3394 nm/% for brands A and B, respectively. Additionally, temperature tests were performed, showing that the sensor works optimally up to 30 °C. The results were validated using a conventional refractometer, showing a close agreement with the data obtained and confirming the reliability and accuracy of the proposed sensor. Compared to other refractometers, the MMI sensor offers advantages such as real-time monitoring, ease of assembly, cost-effectiveness, and minimal maintenance. Furthermore, the sensor represents an alternative tool to guarantee the quality and authenticity of honey, overcoming the limitations of conventional measurement techniques. Full article
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<p>Schematic representation of the SMS fiber optic sensor configuration (sketched by the authors).</p>
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<p>Experimental array to test SMS structure. The inset shows in detail the spectrum of the SMS sensor.</p>
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<p>Spectral response of the SMS sensor, as a reference the spectral response in the initial condition is also represented, that is, when the sensor is surrounded by air. (<b>a</b>) Sensor response for brand A honey mixtures with glucose adulteration. (<b>b</b>) Sensor response for brand A honey mixtures with different moisture contents.</p>
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<p>Spectral shift of the maximum wavelength and refractive index (RI) relative to the initial condition, based on the two types of adulteration analyzed: glucose and moisture content in honey for the two honey brands examined. (<b>a</b>) Measurement of brand A honey adulterated with glucose using the SMS sensor versus a refractometer. (<b>b</b>) Measurement of brand B honey adulterated with glucose using the SMS sensor versus a refractometer. (<b>c</b>) Measurement of brand A honey adulterated with distilled water (moisture content) using the SMS sensor versus a refractometer. (<b>d</b>) Measurement of brand B honey adulterated with distilled water (moisture content) using the SMS sensor versus a refractometer.</p>
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20 pages, 6503 KiB  
Article
Design and Validation of a Monte Carlo Method for the Implementation of Noninvasive Wearable Devices for HbA1c Estimation Considering the Skin Effect
by Tae-Ho Kwon, Shifat Hossain, Mrinmoy Sarker Turja and Ki-Doo Kim
Micromachines 2024, 15(9), 1067; https://doi.org/10.3390/mi15091067 - 24 Aug 2024
Viewed by 1009
Abstract
To diagnose diabetes early or to maintain stable blood glucose levels in diabetics, blood glucose levels should be frequently checked. However, the only way to check blood glucose levels regularly is to use invasive methods, such as pricking the fingertip or using a [...] Read more.
To diagnose diabetes early or to maintain stable blood glucose levels in diabetics, blood glucose levels should be frequently checked. However, the only way to check blood glucose levels regularly is to use invasive methods, such as pricking the fingertip or using a minimally invasive patch. These invasive methods pose several problems, including being painful and potentially causing secondary infections. This study focuses on noninvasively measuring glycated hemoglobin (HbA1c) using PPG signals. In particular, the study relates to a method and a hardware design technology for removing noise that may be present in a PPG signal due to skin contact with a noninvasive HbA1c measurement device. The proposed HbA1c measurement device consists of the first sensor (PPG sensor) module including an optical barrier and the second sensor (cylindrical sensor) module for removing the skin effect. We have developed a Monte Carlo method to implement accurate, noninvasive HbA1c measurement by considering different skin properties among different subjects. Implementing this model in wearable devices will allow end users to not only monitor their glycated hemoglobin levels but also control diabetes with higher accuracy without needing any blood samples. This will be a groundbreaking advancement in modern wearable medical devices. Full article
(This article belongs to the Special Issue The 15th Anniversary of Micromachines)
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<p>PPG signal impurities in a PPG sensor.</p>
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<p>Hardware design configuration of a watch-type PPG sensor for HbA1c estimation.</p>
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<p>(<b>a</b>) LED and PD arrangement in the device and (<b>b</b>) block diagram of device components [<a href="#B11-micromachines-15-01067" class="html-bibr">11</a>].</p>
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<p>PPG data acquisition and corresponding measurement results.</p>
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<p>Peaks and valleys of a subject’s three-wavelength signal to determine AC and DC components (The black line represents the average value of the signal, the purple line represents the average value of the peaks, and the yellow line represents the average value of the valleys.): (<b>a</b>) red wavelength, (<b>b</b>) green wavelength, and (<b>c</b>) blue wavelength.</p>
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<p>(<b>a</b>) AC/DC versus HbA1c for 28 subjects and (<b>b</b>) ratio versus HbA1c for 28 subjects.</p>
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<p>Overall estimation process diagram with the Monte Carlo simulation model.</p>
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<p>The MCS model where photons will traverse through (The red circles represent photons).</p>
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<p>Monte Carlo simulation flow chart.</p>
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<p>(<b>a</b>) Simulated AC/DC (blue) vs. HbA1c, (<b>b</b>) simulated AC/DC (green) vs. HbA1c, and (<b>c</b>) simulated AC/DC (red) vs. HbA1c with variable SpO2 values.</p>
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<p>(<b>a</b>) Simulated R1 vs. HbA1c, (<b>b</b>) simulated R2 vs. HbA1c, and (<b>c</b>) simulated R3 vs. HbA1c with variable SpO2 values.</p>
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<p>XGBoost calibration model.</p>
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<p>EGA and Bland–Altman analysis for HbA1c considering 3 AC/DCs and 3 cylindrical values.</p>
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<p>EGA and Bland–Altman analysis for HbA1c estimation considering 2 AC/DCs, 1 ratio, and 3 cylindrical values.</p>
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<p>(<b>a</b>) AC/DC versus HbA1c for 50 subjects, (<b>b</b>) Ratio versus HbA1c for 50 subjects.</p>
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<p>EGA and Bland–Altman analysis for HbA1c considering 3 AC/DCs and 3 cylindrical values in 50 subjects.</p>
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<p>EGA and Bland–Altman analysis for HbA1c when considering 2 AC/DCs, 1 ratio, and 3 cylindrical values (Method 2) in 50 subjects.</p>
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19 pages, 5276 KiB  
Article
Design and Implementation of a Low-Power Device for Non-Invasive Blood Glucose
by Luis Miguel Pires and José Martins
Designs 2024, 8(4), 63; https://doi.org/10.3390/designs8040063 - 24 Jun 2024
Cited by 1 | Viewed by 1964
Abstract
Glucose is a simple sugar molecule. The chemical formula of this sugar molecule is C6H12O6. This means that the glucose molecule contains six carbon atoms (C), twelve hydrogen atoms (H), and six oxygen atoms (O). In human [...] Read more.
Glucose is a simple sugar molecule. The chemical formula of this sugar molecule is C6H12O6. This means that the glucose molecule contains six carbon atoms (C), twelve hydrogen atoms (H), and six oxygen atoms (O). In human blood, the molecule glucose circulates as blood sugar. Normally, after eating or drinking, our bodies break down the sugars in food and use them to obtain energy for our cells. To execute this process, our pancreas produces insulin. Insulin “pulls” sugar from the blood and puts it into the cells for use. If someone has diabetes, their pancreas cannot produce enough insulin. As a result, the level of glucose in their blood rises. This can lead to many potential complications, including blindness, disease, nerve damage, amputation, stroke, heart attack, damage to blood vessels, etc. In this study, a non-invasive and therefore easily usable method for monitoring blood glucose was developed. With the experiment carried out, it was possible to measure glucose levels continuously, thus eliminating the disadvantages of invasive systems. Near-IR sensors (optical sensors) were used to estimate the concentration of glucose in blood; these sensors have a wavelength of 940 nm. The sensor was placed on a small black parallelepiped-shaped box on the tip of the finger and the output of the optical sensor was then connected to a microcontroller at the analogue input. Another sensor used, but only to provide more medical information, was the heartbeat sensor, inserted into an armband (along with the microprocessor). After processing and linear regression analysis, the glucose level was predicted, and data were sent via the Bluetooth network to a developed APP. The results of the implemented device were compared with available invasive methods (commercial products). The hardware consisted of a microcontroller, a near-IR optical sensor, a heartbeat sensor, and a Bluetooth module. Another objective of this experiment using low-cost and low-power hardware was to not carry out complex processing of data from the sensors. Our practical laboratory experiment resulted in an error of 2.86 per cent when compared to a commercial product, with a hardware cost of EUR 8 and a consumption of 50 mA. Full article
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<p>Light spectrum [<a href="#B23-designs-08-00063" class="html-bibr">23</a>].</p>
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<p>System architecture.</p>
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<p>NIR emitter LED (<b>left</b>) and optical receiver (<b>right</b>).</p>
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<p>MAX30102 board (5-pin version).</p>
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<p>MAX30102 board (7-pin version).</p>
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<p>MAX30102 block diagram.</p>
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<p>Blood circulation example: arteries (<b>a</b>) and veins (<b>b</b>).</p>
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<p>Electrical scheme of experiment.</p>
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<p>Firmware flowchart.</p>
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<p>Influence of light propagation in glucose molecules.</p>
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<p>Experimental diagram of the glucose meter.</p>
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<p>Point cloud and linear regression to expression in (1).</p>
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<p>Visual design of the experimental prototype.</p>
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<p>Appearance of the app.</p>
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13 pages, 3871 KiB  
Article
Detection of NT-proBNP Using Optical Fiber Back-Reflection Plasmonic Biosensors
by Ana Sofia Assunção, Miguel Vidal, Maria João Martins, Ana Violeta Girão, Médéric Loyez, Christophe Caucheteur, José Mesquita-Bastos, Florinda M. Costa, Sónia O. Pereira and Cátia Leitão
Biosensors 2024, 14(4), 173; https://doi.org/10.3390/bios14040173 - 4 Apr 2024
Cited by 3 | Viewed by 4741
Abstract
Heart failure (HF) is a clinical entity included in cardiovascular diseases affecting millions of people worldwide, being a leading cause of hospitalization of older adults, and therefore imposing a substantial economic burden on healthcare systems. HF is characterized by dyspnea, fatigue, and edema [...] Read more.
Heart failure (HF) is a clinical entity included in cardiovascular diseases affecting millions of people worldwide, being a leading cause of hospitalization of older adults, and therefore imposing a substantial economic burden on healthcare systems. HF is characterized by dyspnea, fatigue, and edema associated with elevated blood levels of natriuretic peptides, such as N Terminal pro-B-type Natriuretic Peptide (NT-proBNP), for which there is a high demand for point of care testing (POCT) devices. Optical fiber (OF) biosensors offer a promising solution, capable of real-time detection, quantification, and monitoring of NT-proBNP concentrations in serum, saliva, or urine. In this study, immunosensors based on plasmonic uncladded OF tips were developed using OF with different core diameters (200 and 600 µm). The tips were characterized to bulk refractive index (RI), anddetection tests were conducted with NT-proBNP concentrations varying from 0.01 to 100 ng/mL. The 200 µm sensors showed an average total variation of 3.6 ± 2.5 mRIU, an average sensitivity of 50.5 mRIU/ng·mL−1, and a limit of detection (LOD) of 0.15 ng/mL, while the 600 µm sensors had a response of 6.1 ± 4.2 mRIU, a sensitivity of 102.8 mRIU/ng·mL−1, and an LOD of 0.11 ng/mL. Control tests were performed using interferents such as uric acid, glucose, and creatinine. The results show the potential of these sensors for their use in biological fluids. Full article
(This article belongs to the Special Issue Plasmonic Biosensors for Biomedical Applications)
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<p>Pictures of (<b>a</b>) the gold sputter coater, where OF tips are placed horizontally inside the vacuum chamber, (<b>b</b>) the sputtering process, (<b>c</b>) the batches of 200 µm tips and 600 µm tips, ready to be annealed, and (<b>d</b>–<b>f</b>) SEM images of an OF tip after thermal annealing.</p>
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<p>Schematic diagram of the experimental set-up for a gold coated OF tip (image not to scale).</p>
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<p>Plasmonic OF tip and respective biofunctionalization steps for biosensor fabrication.</p>
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<p>Characterization to RI: reflectivity spectra recorded in solutions with RI ranging from 1.3330 to 1.3853 RIU for (<b>a</b>) a 200 µm tip and (<b>b</b>) a 600 µm tip; (<b>c</b>) <span class="html-italic">λ<sub>SPR</sub></span> shift as a function of RI with respective polynomial fits, Δ<span class="html-italic">λ<sub>SPR</sub></span> = 18,131.3RI<sup>2</sup> − 47,552.7RI + 31,783.3 and Δ<span class="html-italic">λ<sub>SPR</sub></span> = 22,346.7RI<sup>2</sup> − 58,810.3RI + 39,293.1 for the 200 µm tips (n = 13) and 600 µm tips (n = 13), respectively; (<b>d</b>) RI sensitivities expressed as <span class="html-italic">S</span> = 36,262.6RI – 47,552.7 and <span class="html-italic">S</span> = 44,693.3RI − 58,810.3 for the 200 µm tips and 600 µm tips, respectively.</p>
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<p>Tips response during biofunctionalization: reflection spectra acquired in PBS after each functionalization step with respective Δ<span class="html-italic">λ<sub>SPR</sub></span> for (<b>a</b>,<b>b</b>) 200 µm tips and (<b>c</b>,<b>d</b>) 600 µm tips.</p>
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<p>Reflectivity spectra, acquired in PBS, after 30 min incubation periods in NT-proBNP solutions in a concentration range of 0.01–100 ng/mL: (<b>a</b>) 200 µm tip and (<b>b</b>) 600 µm tip.</p>
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<p>Example of a tip sensor response over time, exposing a 200 µm tip biosensor to 5 different concentrations of NT-proBNP, in steps of 30 min, for a total of 230 min (the chart displays the raw data in dots, with the black dots being the acquisitions in NT-proBNP solutions and the red dots in PBS solution after the washing step, and the average trace in blue).</p>
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<p>Tips response in ∆RI of (<b>a</b>) 200 µm tips (n = 3) and (<b>b</b>) 600 µm tips (n = 3), exposed to NT-proBNP concentrations ranging from 0.01 ng/mL to 100 ng/mL, showing the Langmuir–Freundlich fit to the experimental data.</p>
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<p>Bar graph comparison of the shift in resonance wavelength for the negative control (solution of glucose, acid uric, and creatinine) and two positive controls with different concentrations of NT-proBNP diluted in the interferent molecules’ solution, for a 600 µm tip.</p>
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12 pages, 3728 KiB  
Article
A Novel Biosensor for the Detection of Glucose Concentration Using the Dual-Peak Long Period Grating in the Near- to Mid-Infrared
by Namita Sahoo, Bing Sun, Yidong Tan, Kaiming Zhou and Lin Zhang
Sensors 2024, 24(4), 1247; https://doi.org/10.3390/s24041247 - 15 Feb 2024
Cited by 2 | Viewed by 1936
Abstract
In this article, we demonstrate an improved efficient fibre sensor with a high sensitivity to measure glucose concentrations in the physiological range of human beings, operating in a broad spectral bandwidth from the near- to mid-infrared. The sensor consists of a dual-peak long [...] Read more.
In this article, we demonstrate an improved efficient fibre sensor with a high sensitivity to measure glucose concentrations in the physiological range of human beings, operating in a broad spectral bandwidth from the near- to mid-infrared. The sensor consists of a dual-peak long period grating (DPLPG) with a period of 150 μm inscribed in an optical fibre with a diameter of 80 μm. The investigation of sensing for refractive index results in a sensitivity of ~−885.7 nm/refractive index unit (RIU) and ~2008.6 nm/RIU in the range of 1.30–1.44. The glucose measurement is achieved by the immobilisation of a layer of enzyme of glucose oxidase (GOD) onto the fibre surface for the selective enhancement of sensitivity for glucose. The sensor can measure glucose concentrations with a maximum sensitivity of −36.25 nm/(mg/mL) in the range of 0.1–3.0 mg/mL. To the best of our knowledge, this is the highest sensitivity ever achieved for a measurement of glucose with a long period grating-based sensor, indicating its potential for many applications including pharmaceutical, biomedical and food industries. Full article
(This article belongs to the Special Issue Fiber Grating Sensors and Applications)
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<p>Schematic of the cladding mode coupling for a long period grating (LPG).</p>
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<p>Transmission spectra for a DPLPG with 150 μm period UV-inscribed in SM1500 (4.2/80) fibre: (<b>a</b>) spectrum from 1100 nm to 1250 nm showing an individual peak, (<b>b</b>) spectrum from 1200 nm to 2100 nm showing the dual-peak feature of around 1650 nm area.</p>
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<p>Temperature response in the range of 10–50 °C for a dualpeak LPG UV inscribed in the thin cladding SM1500 (4.2/80) fibre with a 150 μm period: (<b>a</b>) spectral evolution for the normal blue shifting peak 1, (<b>b</b>) spectral evolution for the dual-peaks 2 and 3 which shift in the opposite directions, (<b>c</b>) resonance wavelength versus increased temperature for all resonances.</p>
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<p>SRI response in the range of 1.30 to 1.44 for the DPLPG UV inscribed in the thin cladding SM1500 (4.2/80) fibre with a 150 μm period: (<b>a</b>) spectral evolution for the normal blue shifting peak 1, (<b>b</b>) spectral evolution for dual-peaks 2 and 3 which move in opposite directions, (<b>c</b>) SRI sensitivity result for peak 1 with increased SRI, (<b>d</b>) SRI sensitivity results for peak 2 and 3 with increased SRI.</p>
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<p>The functionalisation process for the DPLPG fibre sensor surface with (<b>a</b>) cleaning of the fibre, (<b>b</b>) APTES silanisation and (<b>c</b>) during and after GOD immobilisation.</p>
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<p>Spectral evolution for DPLPG after each process of surface treatment, silanisation with APTES and enzyme functionalisation with GOD for (<b>a</b>) lower wavelength peak 1 and (<b>b</b>) higher wavelength peaks 2 and 3.</p>
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<p>Schematic of the experimental setup for glucose sensing of enzyme-immobilised DPLPG.</p>
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<p>Spectral evaluation for enzyme functionalised DPLPG with various glucose concentrations for (<b>a</b>) resonance peak 1 and (<b>b</b>) resonance peaks 2 and 3. Sensitivity analysis for (<b>c</b>) normal single peak 1 and (<b>d</b>) conjugated attenuation peaks 2 and 3.</p>
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14 pages, 5506 KiB  
Article
Enhanced Sensitivity in Optical Sensors through Self-Image Theory and Graphene Oxide Coating
by Cristina Cunha, Catarina Monteiro, António Vaz, Susana Silva, Orlando Frazão and Susana Novais
Sensors 2024, 24(3), 891; https://doi.org/10.3390/s24030891 - 30 Jan 2024
Cited by 4 | Viewed by 2487
Abstract
This paper presents an approach to enhancing sensitivity in optical sensors by integrating self-image theory and graphene oxide coating. The sensor is specifically engineered to quantitatively assess glucose concentrations in aqueous solutions that simulate the spectrum of glucose levels typically encountered in human [...] Read more.
This paper presents an approach to enhancing sensitivity in optical sensors by integrating self-image theory and graphene oxide coating. The sensor is specifically engineered to quantitatively assess glucose concentrations in aqueous solutions that simulate the spectrum of glucose levels typically encountered in human saliva. Prior to sensor fabrication, the theoretical self-image points were rigorously validated using Multiphysics COMSOL 6.0 software. Subsequently, the sensor was fabricated to a length corresponding to the second self-image point (29.12 mm) and coated with an 80 µm/mL graphene oxide film using the Layer-by-Layer technique. The sensor characterization in refractive index demonstrated a wavelength sensitivity of 200 ± 6 nm/RIU. Comparative evaluations of uncoated and graphene oxide-coated sensors applied to measure glucose in solutions ranging from 25 to 200 mg/dL showed an eightfold sensitivity improvement with one bilayer of Polyethyleneimine/graphene. The final graphene oxide-based sensor exhibited a sensitivity of 10.403 ± 0.004 pm/(mg/dL) and demonstrated stability with a low standard deviation of 0.46 pm/min and a maximum theoretical resolution of 1.90 mg/dL. Full article
(This article belongs to the Special Issue Advanced Optics and Photonics Technologies for Sensing Applications)
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<p>Sensor structure design, where <span class="html-italic">L<sub>CSF</sub></span> corresponds to the CSF length—image not to scale.</p>
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<p>(<b>a</b>) Light propagation; (<b>b</b>) electric field distribution (longitudinal) for a CSF tip with a length of 29.12 mm.</p>
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<p>Scheme of the experimental setup for cleavage.</p>
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<p>Description of the LbL process.</p>
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<p>Spectra results for sensor with 2 bilayers of PEI/GO.</p>
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<p>SEM images of (<b>a</b>,<b>b</b>) uncoated sensors; (<b>c</b>,<b>d</b>) coated sensors with one bilayer of PEI/GO.</p>
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<p>Schematic configuration of the experimental setup.</p>
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<p>Sensor response to the RI experiment.</p>
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<p>Output spectra of the GO-based CSF tip.</p>
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<p>Sensor response to glucose concentration variations.</p>
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<p>Stability test for different sensors coated with one bilayer of PEI/GO.</p>
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25 pages, 5212 KiB  
Review
Application of Optical Fiber Sensing Technology and Coating Technology in Blood Component Detection and Monitoring
by Wenwen Qu, Yanxia Chen, Chaoqun Ma, Donghong Peng, Xuanyao Bai, Jiaxin Zhao, Shuangqiang Liu and Le Luo
Coatings 2024, 14(2), 173; https://doi.org/10.3390/coatings14020173 - 30 Jan 2024
Cited by 2 | Viewed by 3759
Abstract
The advantages of optical fiber sensors include their miniaturization, strong anti-interference ability, high sensitivity, low cost, and fast response speed. They can be used for in situ detection in harsh environments, making them suitable for a wide range of applications such as blood [...] Read more.
The advantages of optical fiber sensors include their miniaturization, strong anti-interference ability, high sensitivity, low cost, and fast response speed. They can be used for in situ detection in harsh environments, making them suitable for a wide range of applications such as blood detection and monitoring. This technology holds great potential for medical diagnosis and health monitoring, opening up new possibilities in the field. Coating technology plays a crucial role in enhancing the sensitivity and stability of optical fiber sensors, ultimately improving their measurement accuracy and reliability. This manuscript expounds the application status and progression of optical fiber sensors in the determination of blood glucose concentrations, blood pH, diverse proteins in blood, and physical properties of blood. The principle of optical fiber sensors and the application of coating technology for detecting varying targets are scrutinized in detail, with particular emphasis on the advantages and limitations of distinct design schemes. The adept amalgamation of optical fiber sensing technology and coating technology amplifies the adaptability of optical fiber sensors in diverse practical scenarios, thereby presenting novel instruments and methodologies for researchers in pertinent fields to augment their advancement and development. Full article
(This article belongs to the Special Issue Optical Coatings: From Materials to Applications)
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<p>Schematic diagram of the principle of an optical fiber sensor. OSA: optic spectral analyzer.</p>
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<p>Experimental setup for glucose detection. The upward inset shows the schematical dissection diagram of GOD-GO-modified TFBG. BBS: broadband source; PC: polarization controller [<a href="#B51-coatings-14-00173" class="html-bibr">51</a>].</p>
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<p>Schematic diagram of GOD immobilization onto a multimode microfiber. (<b>a</b>) Hydroxyl-groups-activated microfiber; (<b>b</b>) APTES-coated microfiber; (<b>c</b>) GOD-immobilized microfiber [<a href="#B66-coatings-14-00173" class="html-bibr">66</a>].</p>
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<p>(<b>a</b>) Experimental setup of the sensing system; (<b>b</b>) Schematic of the hydrogel-coating optical fiber SPR sensing part [<a href="#B103-coatings-14-00173" class="html-bibr">103</a>].</p>
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<p>(<b>a</b>) Schematic diagram of the proposed MAMZI structure; (<b>b</b>) Microscopic image of the fabricated MAMZI; (<b>c</b>) Photograph of the encapsulated microfluidic chip [<a href="#B109-coatings-14-00173" class="html-bibr">109</a>].</p>
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<p>(<b>a</b>) Cr/Au coating on a fiber core; (<b>b</b>) Asymmetric cross-section of metal layers coated on the fiber core [<a href="#B135-coatings-14-00173" class="html-bibr">135</a>].</p>
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<p>The schematic overview of the FBG sensor [<a href="#B139-coatings-14-00173" class="html-bibr">139</a>].</p>
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<p>(<b>a</b>) Plan view of the probe; (<b>b</b>) Side view of the probe [<a href="#B158-coatings-14-00173" class="html-bibr">158</a>].</p>
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32 pages, 7515 KiB  
Review
MOF-Based Materials for Glucose Detection
by Yiling Zhang, Qian Lin, Yiteng Song, Jiaqi Huang, Miaomiao Chen, Runqi Ouyang, Si-Yang Liu and Zong Dai
Chemosensors 2023, 11(8), 429; https://doi.org/10.3390/chemosensors11080429 - 2 Aug 2023
Cited by 15 | Viewed by 4602
Abstract
Metal–organic frameworks (MOFs), constructed by coordination between metal-containing nodes and organic linkers, are widely used in various fields due to the advantages of tunable pores, diverse functional sites, stable structure, and multi-functionality. It should be noted that MOF-based materials play a major role [...] Read more.
Metal–organic frameworks (MOFs), constructed by coordination between metal-containing nodes and organic linkers, are widely used in various fields due to the advantages of tunable pores, diverse functional sites, stable structure, and multi-functionality. It should be noted that MOF-based materials play a major role in glucose detection, serving as a signal transducer or functional substrate for embedding nanoparticles/enzymes. Diabetes is one of the most common and fast-growing diseases worldwide, whose main clinical manifestation is high blood sugar levels. Therefore, accurate, sensitive, and point-of-care glucose detection is necessary. This review orderly introduces general synthetic strategies of MOF-based materials (pristine MOF, nanoparticles, or enzymes-modified MOF and MOF-derived materials) and detection methods (electrochemical and optical methods) for glucose detection. Then, the review refers to the novel MOF-based glucose detection devices (flexible wearable devices and microfluidic chips), which enable non-invasive continuous glucose monitoring or low-cost microscale detection. On the basis of describing the development of glucose sensors based on MOF materials in the past five years, the review presents merits, demerits, and possible improvements of various detection methods. Full article
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<p>Different types of MOF-based materials for glucose detection. Pristine MOFs: Cu-BTC (benzene-1,3,5-tricarboxylate) (reproduced with permission from [<a href="#B15-chemosensors-11-00429" class="html-bibr">15</a>], copyright © 2013, Royal Society of Chemistry), ZIF-67/8 (reproduced with permission from [<a href="#B16-chemosensors-11-00429" class="html-bibr">16</a>], copyright © 2022, Elsevier); NPs modified MOFs: AgNPs@ZIF-67 (reproduced with permission from [<a href="#B17-chemosensors-11-00429" class="html-bibr">17</a>], copyright © 2018, Elsevier), Pt/Fe-MOF (reproduced with permission from [<a href="#B18-chemosensors-11-00429" class="html-bibr">18</a>], copyright © 2021, Springer Nature); Enzymes modified MOFs: GOx&amp;Luminol@ZIF-67 (reproduced with permission from [<a href="#B19-chemosensors-11-00429" class="html-bibr">19</a>], copyright © 2023, Elsevier); MOF-derived materials: Ni/NiO (reproduced with permission from [<a href="#B20-chemosensors-11-00429" class="html-bibr">20</a>], copyright © 2020, Springer Nature), ST-Co<sub>3</sub>O<sub>4</sub> (reproduced with permission from [<a href="#B21-chemosensors-11-00429" class="html-bibr">21</a>], copyright © 2021, John Wiley and Sons). The typical works of electrochemical methods (reproduced with permission from [<a href="#B22-chemosensors-11-00429" class="html-bibr">22</a>], copyright © 2018, Elsevier) and optical methods (reproduced with permission from [<a href="#B23-chemosensors-11-00429" class="html-bibr">23</a>], copyright © 2019, John Wiley and Sons) for glucose detection with MOF-based materials. Furthermore, the flexible wearable sensors (reproduced with permission from [<a href="#B24-chemosensors-11-00429" class="html-bibr">24</a>], copyright © 2022, Elsevier) and microfluidic chips (reproduced with permission from [<a href="#B25-chemosensors-11-00429" class="html-bibr">25</a>], copyright © 2022, Elsevier) for MOF-based glucose detection.</p>
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<p>Synthesis methods of the pristine MOF. (<b>A</b>) Solvothermal synthesis of NiCo-BTC/CC (reproduced with permission from [<a href="#B39-chemosensors-11-00429" class="html-bibr">39</a>], copyright © 2022, American Chemical Society). (<b>B</b>) Room-temperature synthesis of 2D Fe-BTC (reproduced with permission from [<a href="#B41-chemosensors-11-00429" class="html-bibr">41</a>], copyright © 2013, Royal Society of Chemistry). (<b>C</b>) Microwave-assisted synthesis of ZIF-67/8 (reproduced with permission from [<a href="#B16-chemosensors-11-00429" class="html-bibr">16</a>], copyright © 2022, Elsevier). (<b>D</b>) Electrochemical deposition synthesis of Cu-BTC on SWCNTs/GCE (reproduced with permission from [<a href="#B42-chemosensors-11-00429" class="html-bibr">42</a>], copyright © 2020, Elsevier).</p>
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<p>The synthesis of NPs-modified MOFs. NPs coated on MOFs: (<b>A</b>) AuNPs/N-GQDs coated on P-MOF as a substrate (reproduced with permission from [<a href="#B66-chemosensors-11-00429" class="html-bibr">66</a>], copyright © 2022, Elsevier); (<b>B</b>) PtNPs coated on Fe-MOF as a catalyst (reproduced with permission from [<a href="#B18-chemosensors-11-00429" class="html-bibr">18</a>], copyright © 2021, Springer Nature). NPs encapsulated in MOFs: (<b>C</b>) BODIPY encapsulated in Eu-MOF by one-pot synthesis (reproduced with permission from [<a href="#B67-chemosensors-11-00429" class="html-bibr">67</a>], copyright © 2022, Elsevier); (<b>D</b>) AgNPs encapsulated in Co-MOF by in situ growth (reproduced with permission from [<a href="#B68-chemosensors-11-00429" class="html-bibr">68</a>], copyright © 2019, American Chemical Society).</p>
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<p>The synthesis of enzymes-modified MOFs. Enzymes coated on MOFs: (<b>A</b>) GOx-Fe<sub>3</sub>Ni-MOF by physical adsorption (reproduced with permission from [<a href="#B79-chemosensors-11-00429" class="html-bibr">79</a>], copyright © 2022, American Chemical Society); (<b>B</b>) GOx&amp;Hemin-ZIF-67 by the coordination between the Co<sup>2+</sup> and the carbonyl group of the proteins (reproduced with permission from [<a href="#B24-chemosensors-11-00429" class="html-bibr">24</a>], copyright © 2022, Elsevier). Enzymes encapsulated in MOFs: (<b>C</b>) HRP-PAA@ZIF-L (reproduced with permission from [<a href="#B80-chemosensors-11-00429" class="html-bibr">80</a>], copyright © 2022, Elsevier); (<b>D</b>) GOx@FCM (reproduced with permission from [<a href="#B81-chemosensors-11-00429" class="html-bibr">81</a>], copyright © 2012, Royal Society of Chemistry).</p>
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<p>(<b>A</b>) The synthesis strategy of CuO/C (reproduced with permission from [<a href="#B98-chemosensors-11-00429" class="html-bibr">98</a>], copyright © 2022, Springer Nature). (<b>B</b>) The synthesis strategy of YASNiCo@C (reproduced with permission from [<a href="#B99-chemosensors-11-00429" class="html-bibr">99</a>], copyright © 2020, Royal Society of Chemistry).</p>
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<p>(<b>A</b>) Mechanism of glucose oxidation at swnt-MOF(Cu)@gwe and corresponding CA curves for glucose detection (reproduced with permission from [<a href="#B119-chemosensors-11-00429" class="html-bibr">119</a>], copyright © 2023, Elsevier). (<b>B</b>) Mechanism of glucose electrocatalytic oxidation at Ni@Cu-MOF nanocomposite and corresponding CV curves of glucose detection (reproduced with permission from [<a href="#B120-chemosensors-11-00429" class="html-bibr">120</a>], copyright © 2020, Elsevier). (<b>C</b>) Fabrication of NPC-Co<sub>3</sub>O<sub>4</sub> composite and corresponding DPV curves for glucose detection (reproduced with permission from [<a href="#B22-chemosensors-11-00429" class="html-bibr">22</a>], copyright © 2018, Elsevier). (<b>D</b>) Synthesis of core-shell UiO-67@Ni-MOF composites and corresponding AMP curves for glucose detection (reproduced with permission from [<a href="#B121-chemosensors-11-00429" class="html-bibr">121</a>], copyright © 2020, Elsevier).</p>
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<p>(<b>A</b>) Schematic diagram of the synthesis of Fe<sub>3</sub>Ni-MOF/GOx and the reaction mechanism for the cascade oxidation of glucose (reproduced with permission from [<a href="#B79-chemosensors-11-00429" class="html-bibr">79</a>], copyright © 2022, American Chemical Society). (<b>B</b>) Schematic representation of the preparation of dZIF-8 BH and the colorimetric sensing mechanism based on the biocatalytic cascade of dZIFs BH (reproduced with permission from [<a href="#B91-chemosensors-11-00429" class="html-bibr">91</a>], copyright © 2022, American Chemical Society). (<b>C</b>) Schematic illustration of the principle of glucose detection by G&amp;L@ZIF@Paper and the step-by-step flow for detecting glucose in saliva (reproduced with permission from [<a href="#B19-chemosensors-11-00429" class="html-bibr">19</a>], copyright © 2023, Elsevier).</p>
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<p>(<b>A</b>) Reaction mechanism of FL detection of glucose based on an Eu-MOF (reproduced with permission from [<a href="#B157-chemosensors-11-00429" class="html-bibr">157</a>], copyright © 2019, Elsevier). (<b>B</b>) The smartphone-based Ag NPs/UiO-66-NH<sub>2</sub> and OPD composite film for glucose detection (reproduced with permission from [<a href="#B160-chemosensors-11-00429" class="html-bibr">160</a>], copyright © 2021, American Chemical Society). (<b>C</b>) Synthesis of the GOx@MAF-2 composite and its application in glucose detection (reproduced with permission from [<a href="#B23-chemosensors-11-00429" class="html-bibr">23</a>], copyright © 2019, John Wiley and Sons). (<b>D</b>) CL enhancement mechanism of the luminol-H<sub>2</sub>O<sub>2</sub> system by MOF-235/<span class="html-italic">β</span>-CD composite (reproduced with permission from [<a href="#B23-chemosensors-11-00429" class="html-bibr">23</a>], copyright © 2018, Elsevier). (<b>E</b>) Principle of the novel CL sensor for one-step ultrasensitive glucose detection (i) preparation of Co-TCPP(Fe)@luminol@GOD; (ii) process of one-step detection for glucose (reproduced with permission from [<a href="#B90-chemosensors-11-00429" class="html-bibr">90</a>], copyright © 2019, American Chemical Society). (<b>F</b>) Fabrication and detection principle of MMA (reproduced with permission from [<a href="#B159-chemosensors-11-00429" class="html-bibr">159</a>], copyright © 2022, Springer Nature).</p>
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<p>(<b>A</b>) Schematic illustration of (i) production of rGO/PU fiber, (ii) manufacturing process of the NCGP working electrode and Ag/AgCl fiber reference electrode, (iii) the NCGP glucose sensor integrated into the elastic fabric, and (iv) the physical image of the device attached to the volunteer’s arm (reproduced with permission from [<a href="#B140-chemosensors-11-00429" class="html-bibr">140</a>], copyright © 2021, American Chemical Society). (<b>B</b>) The fabrication process of biomimetic Murray Cu<sub>2</sub>(NDC)<sub>2</sub>/PDHP (reproduced with permission from [<a href="#B148-chemosensors-11-00429" class="html-bibr">148</a>], copyright © 2018, John Wiley and Sons). (<b>C</b>) (i) Schematic illustration of the synthetic route of GOx/Hemin@NC-ZIF; (ii) Left: A photograph of a sweatband integrated with the portable prototype glucose sensor. Middle: Scheme of the GOx/Hemin@NC-ZIF catalytic cascade reaction based on the electrochemical biosensor and the all-integrated sensor fabricated on a polyimide (PI) sheet. Right: A photograph of the smartphone with an app for the perspiration analysis (reproduced with permission from [<a href="#B24-chemosensors-11-00429" class="html-bibr">24</a>], copyright © 2022, Elsevier).</p>
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<p>(<b>A</b>) (i) Schematic of the well-based μPAD; (ii) well-based μPAD layers; and (iii) assembled chip. (<b>B</b>) (i) Schematic of LFA-based μPAD; (ii) layers; and (iii) assembled chip (reproduced with permission from [<a href="#B162-chemosensors-11-00429" class="html-bibr">162</a>], copyright © 2019, Elsevier). (<b>C</b>) Schematic design for the simultaneous determination of different sugars using a CeO<sub>2</sub>@NH<sub>2</sub>-MIL-88B(Fe)-modified paper-based analytical device (reproduced with permission from [<a href="#B25-chemosensors-11-00429" class="html-bibr">25</a>], copyright © 2022, Elsevier).</p>
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14 pages, 2459 KiB  
Article
Glucose Sensor Using Sol–Gel Coating Layer Deposited on PMMA Optical Fiber: An Enzyme Activity Measurement System
by Jorge-A. Ortega-Contreras, Edgar Alvarado-Méndez, Guillermo Almanza-Rodríguez, María del Carmen Hernández and Luis Celaya-García
Gels 2023, 9(8), 608; https://doi.org/10.3390/gels9080608 - 27 Jul 2023
Cited by 1 | Viewed by 1608
Abstract
In the present work, a biocatalytic glucose optical sensor produced by immobilizing glucose oxidase (GOD) as a recognition molecule over a PMMA (polymethylmethacrylate) optical fiber is introduced. An enzymatic encapsulation process was carried out using the sol–gel method, depositing a TEOS-based coating by [...] Read more.
In the present work, a biocatalytic glucose optical sensor produced by immobilizing glucose oxidase (GOD) as a recognition molecule over a PMMA (polymethylmethacrylate) optical fiber is introduced. An enzymatic encapsulation process was carried out using the sol–gel method, depositing a TEOS-based coating by immersion at the end of an optical fiber; the biosensor was characterized using different glucose levels. Finally, the best way to encapsulate the enzyme and prevent it from degrading is to perform the process at room temperature, and later implement the deposition of the coating on the fiber. The drying process was optimal below 8 °C. Full article
(This article belongs to the Special Issue Advances and Current Applications in Gel-Based Membranes)
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<p>Bioptrode observed under 50x objective. (<b>a</b>) Front view (<b>b</b>) Side view.</p>
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<p>Bioptrodes.</p>
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<p>Effect of the GOD/HRP staining reaction with o-Dianisidine on power.</p>
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<p>Optical arrangement for colorimetry assay.</p>
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<p>Absorbance in the colorimetry assay.</p>
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<p>HRP enzymatic activity.</p>
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<p>Colorimetry assay tests. (<b>a</b>) Saturation color scale from left to right 10 mM to 0 mM. (<b>b</b>) Power measurement at the highest concentration.</p>
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<p>Schematic of experimental setup in transmission mode to test the bioptrode.</p>
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<p>Normalized intensity at the output of the POF.</p>
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<p>Statistical analysis of the sensitivity of the biosensors’ intensity to exposure to different analyte concentration levels.</p>
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<p>Linear regression.</p>
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<p>Schematic of the dip-coating process. Adapted from [<a href="#B29-gels-09-00608" class="html-bibr">29</a>].</p>
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10 pages, 1790 KiB  
Article
Sugar Detection in Aqueous Solution Using an SMS Fiber Device
by Nailea Mar-Abundis, Yadira Aracely Fuentes-Rubio, René Fernando Domínguez-Cruz and José Rafael Guzmán-Sepúlveda
Sensors 2023, 23(14), 6289; https://doi.org/10.3390/s23146289 - 11 Jul 2023
Cited by 5 | Viewed by 2001
Abstract
We report on the fabrication and testing of a fiber optics sensor based on multimodal interference effects, which aims at the detection of different types of sweeteners dissolved in water. The device, which has a simple structure, commonly known as the SMS configuration, [...] Read more.
We report on the fabrication and testing of a fiber optics sensor based on multimodal interference effects, which aims at the detection of different types of sweeteners dissolved in water. The device, which has a simple structure, commonly known as the SMS configuration, is built by splicing a segment of commercial-grade, coreless multimode fiber (NC-MMF) between two standard single-mode fibers (SMFs). In this configuration, the evanescent field traveling outside the core of the NC-MMF allows the sensing of the refractive index of the surrounding media, making it possible to detect different levels of sugar concentration. The optical sensor was tested with aqueous solutions of glucose, fructose, and sucrose in the concentration range from 0 wt% to 20 wt% at room temperature. The proposed device exhibits a linear response with a sensitivity of 0.1835 nm/wt% for sucrose, 0.1687 nm/wt% for fructose, and 0.1694 nm/wt% for glucose, respectively, with a sensing resolution of around 0.5 wt%. Finally, we show that, despite having similar concentration behavior, some degree of discrimination between the different sugars can be achieved by assessing their thermo-optical response. Full article
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<p>(<b>a</b>) Experimental array to test SMS device for the measurement of sugar concentrations. (<b>b</b>) Typical structure used to perform the multimodal interference effect (MMI) in optical fibers.</p>
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<p>Effective refractive index of aqueous solutions of sugars as a function of the sugar concentration. The plot was made by using the numerical data reported in Ref. [<a href="#B42-sensors-23-06289" class="html-bibr">42</a>].</p>
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<p>(<b>a</b>) Spectral response of the MMI sensor for water–sucrose blends. (<b>b</b>) The shaded box shows in more detail the region of the spectral change in panel (<b>a</b>), which corresponds to the NC-MMF section surrounded by each sugar type concentration. (<b>c</b>) The spectral shift of the peak wavelength, with respect to the baseline condition, as a function of the concentration of sucrose, fructose, and glucose present in the aqueous solution.</p>
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<p>The thermal effect reported for each sugar type using a concentration of 18.5%. Measurements were performed in the temperature range from 25 °C to 42.5 °C in increments of 2.5 °C.</p>
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24 pages, 2585 KiB  
Review
Photoplethysmography in Wearable Devices: A Comprehensive Review of Technological Advances, Current Challenges, and Future Directions
by Kwang Bok Kim and Hyun Jae Baek
Electronics 2023, 12(13), 2923; https://doi.org/10.3390/electronics12132923 - 3 Jul 2023
Cited by 46 | Viewed by 22162
Abstract
Photoplethysmography (PPG) is an affordable and straightforward optical technique used to detect changes in blood volume within tissue microvascular beds. PPG technology has found widespread application in commercial medical devices, enabling measurements of oxygen saturation, blood pressure, and cardiac output; the assessment of [...] Read more.
Photoplethysmography (PPG) is an affordable and straightforward optical technique used to detect changes in blood volume within tissue microvascular beds. PPG technology has found widespread application in commercial medical devices, enabling measurements of oxygen saturation, blood pressure, and cardiac output; the assessment of autonomic nerve function; and the diagnosis of peripheral vascular disease. Recently, the growing demand for non-invasive, portable, cost-effective technology, along with advancements in small semiconductor components, has led to the integration of PPG into various wrist-worn wearable devices. Multiple sensor structures have been proposed and, through appropriate signal processing and algorithmic application, these wearable devices can measure a range of health indicators during daily life. This paper begins by addressing the market status of wrist-worn wearable devices, followed by an explanation of the fundamental principles underlying light operation and its interaction with living tissue for PPG measurements. Moving on to technological advancements, the paper addresses the analog front end for the measurement of the PPG signal, sensor configurations with multiple light emitters and receivers, the minimum sampling rate required for low-power systems, and the measurement of stress, sleep, blood pressure, blood glucose, and activity using PPG signals. Several challenges in the field are also identified, including selecting the appropriate wavelength for the PPG sensor’s light source, developing low-power interpolation methods to extract high-resolution inter-beat intervals at a low sampling rate, and exploring the measurement of physiological phenomena using multi-wavelength PPG signals simultaneously collected at the same location. Lastly, the paper presents future research directions, which encompass the development of new, reliable parameters specific to wearable PPG devices and conducting studies in real-world scenarios, such as 24-h long-term measurements. Full article
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<p>(<b>a</b>) Absorbed and reflected light in living tissue for PPG model. (<b>b</b>) the measurement model of reflectance type PPG through the skin microvascular layer, with emphasis on the role of wavelength. The amount of absorbed light correlates with the pulsation of arterial blood. In the systolic phase, the diameter of the arterial vessels is maximal and therefore the absorbance due to arterial hemoglobin is also maximal and the amount of detected light is low, which corresponds to a sensor peak.</p>
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<p>Absorbed and reflected light in living tissue for PPG model. The amount of absorbed light correlates with the pulsation of arterial blood. In the systolic phase, the diameter of the arterial vessels is maximal and therefore the absorbance due to arterial hemoglobin is also maximal and the amount of detected light is low, which is displayed as a peak.</p>
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<p>Conceptual diagram of a PPG sensor module with an embedded LED–PD pair. The diagram illustrates two scenarios: (<b>a</b>) when the integrated sensor module is in direct contact with the wrist and (<b>b</b>) when the sensor module is embedded in a device and covered with a glass surface. In the diagram, the red solid line represents incident light from the LED entering skin tissue, while the green solid line indicates light causing various reactions inside the tissue. The blue solid line represents light entering the PD from skin tissue and being detected. Additionally, the red dotted line in the diagram illustrates light reflected inside the sensor module that does not enter the skin, some of which directly enters the light-receiving unit. The blue dotted line represents the light that is reflected back by the cover glass among the light entering towards the PD.</p>
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<p>Representative examples of PPG sensor systems applied to commercial wrist-worn wearable devices: (<b>a</b>) the Samsung Galaxy Gear Fit, released in 2014, utilizes a sensor module with an embedded LED–PD optopair, and (<b>b</b>) the Samsung Galaxy Watch 3, released in 2020, uses separate LEDs and PDs as a sensor. An analog front end (AFE) is employed in both devices to measure the PD output as a PPG signal.</p>
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<p>Examples of PPG sensor configurations that use multiple LEDs and PDs, and their application to commercial smartwatches. (<b>a</b>–<b>d</b>) illustrate different PPG sensor structures, while (<b>e</b>) through (<b>h</b>) present examples of these structures in use: (<b>e</b>) structure (<b>a</b>) applied to the Samsung Galaxy Gear s3; (<b>f</b>) structure (<b>d</b>) applied to the Samsung Galaxy Watch; (<b>g</b>) shows the sensor structure of the Garmin Fenix 5 with three LEDs and one PD in the center, and (<b>h</b>) shows the sensor structure of the Apple Watch with two LEDs and two PDs.</p>
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<p>Examples of the low sampling rate problem in pulse rate variability analysis using PPG signals. The gray background line represents the PPG signal at 250 Hz, while the circular marker line represents the PPG signal at a lower sampling rate. The dotted line shows the parabolic approximation, which was estimated using the three largest samples in each pulse of the low-sampling-rate signal. Specifically, the 250-Hz PPG signal was down-sampled to (<b>a</b>) 30 Hz, (<b>c</b>) 20 Hz, and (<b>e</b>)10 Hz, respectively. (<b>b</b>,<b>d</b>,<b>f</b>) are enlarged view of the second peak area of (<b>a</b>,<b>c</b>,<b>e</b>).</p>
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