Biocompatible Materials for Continuous Glucose Monitoring Devices

Electrical Transducers

John X.J. Zhang , Kazunori Hoshino , in Molecular Sensors and Nanodevices, 2014

Capacitive Sensors

A capacitive sensor is based on measurement of changes in capacitance. In practice, sensor designs are very similar to those of conductometric sensors. Figure 4.12 shows a schematic of the sensing principle of a typical capacitive chemical sensor. Absorption of target molecules induces two relevant effects of changes in dielectric constant and swelling. Interdigitated electrodes as described for conductometric sensors are commonly used [5]. Usually measured for sensing is the impedance of the system [6], which includes the resistance of the sensitive layer as shown in the simplified circuit in Fig. 4.12(b). Capacitive sensors may be categorized as a special type of conductometric sensor. The impedance of the sensor is typically found indirectly by incorporating the sensor into an RC circuit. Changes in the resonant frequency of an oscillator, or the level of coupling (or attenuation) of an AC signal is used for measurement.

Figure 4.12. Capacitive Chemical Sensors. (a) Schematic Showing the Working Principle of a Capacitive Chemical Sensor (from [5]). (b) Simplified Equivalent Circuit.

From [6].

Glucose Sensors

Glucose sensors [7–9] are used to measure the blood glucose concentration of a patient and are an important part of managing diabetes mellitus. Type 1 and type 2 diabetes are the most common forms of diabetes. Type 1 diabetes is usually diagnosed in children and young adults and accounts for about 5% of all diagnosed cases of diabetes. Type 2 diabetes has been diagnosed in millions of Americans. According to diabetes report card 2012 issued by National Center for Chronic Disease Prevention and Health Promotion, 18.9% of US adults over 65 years old are diagnosed as diabetes in 2007–2009.

Patients with Type 1 diabetes may test their blood sugar five to ten times a day in order for them to effectively monitor their blood sugar levels. Type 2 diabetics may also consider monitoring their blood sugar levels daily based on their risk for future health complications due to the disease. Blood glucose testing may be also needed for patients with other diseases which may affect the pancreas such as cystic fibrosis. In sports medicine, it is used to monitor physical conditions of athletes. Normal blood glucose levels range between 80–120   mg/dL with spikes reaching up to 250   mg/dL after meals. The sensor should also be able to measure the extremes in blood sugar levels (between 20–500   mg/dL, or 1-30   mM) which a patient may experience during an episode of hyper or hypoglycemia and should have a resolution of ~1   mg/dL, or ~50   µM.

The majority of blood glucose sensors, or glucose meters, are categorized as amperometric sensors, which will be described in this chapter. In chapter 5, we will discuss techniques based on optical transduction such as absorption spectroscopy (see Section 5.5), light scattering and Raman spectroscopy (see Section 5.7). In chapter 7, we describe examples of implantable glucose sensors (see Section 7.3.6).

In amperometric glucose sensors, reducing property of glucose is measured as a current. Sensors contain electrodes to measure the current generated by an enzymatic reaction usually between glucose, an enzyme, and a mediator. Use of glucose oxidase (GOx or GOD) has become the gold standard for glucose sensing [10,11]. The initial concept of glucose enzyme electrodes, where a thin layer of GOx was entrapped via a semipermeable membrane, was introduced by Clark and Lyons [12]. Sensing was based on the measurement of the oxygen consumed by the enzyme-catalyzed reaction

$\begin{array}{c}\mathrm{Glucose}+{\mathrm{O}}_2{\displaystyle \stackrel{{\mathrm{GO}}_{\mathrm{x}}}{⟹}}\mathrm{Gluconic}\mathrm{acid}+{\mathrm{H}}_2{\mathrm{O}}_2\\ \mathrm{Glucose}+{\mathrm{GOx}}_{(\text{ox})}\to \mathrm{Gluconic}\mathrm{acid}+{\mathrm{GOx}}_{(\mathrm{red})}\\ {\mathrm{GOX}}_{(\text{red})}+2{\mathrm{M}}_{(\mathrm{ox})}\to {\mathrm{GOx}}_{(\mathrm{ox})}+2{\mathrm{M}}_{(\mathrm{red})}+2{\mathrm{H}}^{+}\\ 2{\mathrm{M}}_{(\mathrm{red})}\to 2{\mathrm{M}}_{(\mathrm{ox})}+2{\mathrm{e}}^{-}\end{array}$

In this method, glucose reacts with the enzyme GOx_(ox). The reduced enzyme GOx_(red) then reduces two mediator M_(ox) ions to M_(red), which is oxidized back to M_(ox) at the electrode surface. The oxidation process $2{\mathrm{M}}_{(\mathrm{red})}\to 2{\mathrm{M}}_{(\mathrm{ox})}+2{\mathrm{e}}^{-}$ is measured as the current by the electrode. However, for this type of early glucose biosensors, a high operation potential is required to perform the amperometric measurement of hydrogen peroxide. Improved methods utilize artificial mediators instead of oxygen to transfer electrons between the GOx and the electrode [8]. Reduced mediators are formed and reoxidized at the electrode, providing an electrical signal to be measured.

A blood glucose test is typically performed by pricking the finger to draw blood, which is then applied to a disposable "test-strip". Figure 4.13 shows a typical glucose meter and a test strip. Each strip includes layers of electrodes, spacers, immobilized enzymes assembled in a small package. Continued research and development have worked to reduce the overall size of the sensor itself and reduce the amount of blood required for an accurate measurement (~µL).

Figure 4.13. Blood Glucose Sensor. (a) Example of a Commercial Product. (b) Composition of a Test Strip Which Includes Electrodes.

(a) From https://www.accu-chek.com/us/; (b) from [13]. Courtesy of Roche.

The advanced glucose electrodes do not use mediators and measures direct transfer between the enzyme and the electrode. The electrode directly transfers electrons using organic conducting materials based on charge-transfer complexes. This type of electrodes have led to needle-type implantable sensors for continuous blood glucose monitoring [89, 90].

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Sensor Materials, Technologies and Applications

A.N.M. Karim , ... S. Begum , in Comprehensive Materials Processing, 2014

13.20.2.1.1.1 CMOS-MEMS glucose sensor

This glucose sensor applies the working principle of first-generation glucose sensors by detecting the concentration of glucose oxidase (21). However, instead of measuring it through electrochemical methods, the device uses capacitive sensing to detect changes in the dielectric constant of the material. The capacitive sensor is formed using interdigitated composite gold and oxide electrodes placed on a silicon substrate, as shown in Figure 2. Based on eqn [1], the glucose oxidase enzyme converts glucose and oxygen into gluconolactone and hydrogen peroxide. Oxygen is derived from water surrounding the sensor. When glucose is oxidized, the concentration of water decreases and the concentration of hydrogen peroxide increases, resulting in a change in the dielectric constant of the area between the sensing electrodes.

Figure 2. (a) Schematic of integrated glucose sensor. (b) Cross-section and layers of the CMOS-MEMS glucose sensor.

Reproduced from Yang, M. Z.; Dai, C. L.; Hung, C. B. Fabrication of a Glucose Sensor with Oscillator Circuit Using CMOS-MEMS Technique. Microelectron. Eng. September 2012, 97, 353–356.

The dielectric constant of water and hydrogen peroxide is 78 and 60, respectively. Measurement of the fluctuation of dielectric constants can be detected by observing the change in capacitance of the sensor. As shown in Figure 2, the sensing capacitors are connected to a series of odd-numbered inverters such that they form an oscillator circuit. The oscillator circuit generates a frequency output ranging from 17 to 25 MHz, depending on the sensor's capacitance.

Using this simple circuit, the change in capacitance can now be easily monitored by observing the oscillator's output since the frequency change is proportional to the change in capacitance and thus the change in glucose concentration. The sensitivity of the glucose sensor was about 1.48 MHz mM^{− 1}.

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Textured and Porous Materials

Heidi E. Koschwanez , William M. Reichert , in Biomaterials Science (Third Edition), 2013

Porous Coatings to Improve Glucose and Oxygen Transport to Implanted Sensors

Percutaneous glucose sensors must be removed after three to seven days to prevent host inflammation, wound healing, and subsequent foreign-body encapsulation from jeopardizing sensor reliability. The foreign-body response ultimately causes impedance of glucose and oxygen transport to the sensor, resulting in sensor signal deterioration, and frequently sensor failure. Methods that improve analyte (glucose, oxygen) transport to indwelling sensors could allow sensors to reliably measure interstitial glucose concentrations for several weeks, as opposed to days.

Attempts to improve long-term sensor performance have included surface chemical modification, various coatings (Gifford et al., 2005; Nablo et al., 2005; Shin and Schoenfisch, 2006), release of molecular mediators (Friedl, 2004; Ward et al., 2004; Norton et al., 2005), and surface topography (Wisniewski and Reichert, 2000; Wisniewski et al., 2000). The effect of surface texturing on the tissue that surrounds implanted biomaterials is well-documented for devices such as total joint arthroplasty (Bauer and Schils, 1999; Ryan et al., 2006) and percutaneous devices (Tagusari et al., 1998; Walboomers and Jansen, 2005; Kim et al., 2006).

Topographical approaches for improving long-term sensor performance were first proposed by Woodward (1982), who suggested that the best coating for an implanted glucose sensor was a sponge that encourages tissue ingrowth and disrupts fibrosis (Figure I.2.15.5). Efforts to create tissue-modifying textured coatings for implantable sensors are attractive, because their impact is not dependent on a depletable drug reservoir, unlike drug eluting techniques.

FIGURE I.2.15.5. Example of a: (a) Medtronic MiniMed SOF-SENSOR™ glucose sensor; and (b) an experimental porous poly-L-lactic acid (PLLA) coating applied to the sensor tip for investigational purposes. Inset: environmental scanning electron microscope image of porous PLLA coating fabricated using salt-leaching/gas foaming with ammonium bicarbonate (NH₄HCO₃).

(Koschwanez, H. E. (2006). Unpublished data.) Courtesy of John Wiley and Sons.

A significant range of materials and pore sizes are capable of promoting angiogenesis and reducing capsule thickness (Ward et al., 2002). Geometry, rather than chemical composition, of the material appears to determine biomaterial–microvasculature interactions (Brauker et al., 1995; Sieminski and Gooch, 2000). Table I.2.15.3 summarizes leading research in the area of porosity and porous coatings for glucose sensors, including the pore size reported to yield the greatest vascularization and/or the least capsule formation around the implant. Variations in pore size and pore structure in implanted biomaterials may, however, limit the conclusions that can be drawn about how pore size influences tissue response (Marshall et al., 2004).

TABLE I.2.15.3. Summary of Pore Sizes that Yielded Optimal Tissue Response (Promoted Angiogenesis and/or Reduced Capsule Thickness) Around Biomaterials or Glucose Sensors

Investigator	Porous Material	Optimal Pore Size	Application and Test Subject	Duration of Investigation
Brauker et al., 1995	PTFE	5 μm	Membrane implanted in rat subcutis	3 weeks
Sharkawy et al., 1998	PVA	60 μm	Membrane implanted in rat subcutis	12-16 weeks
Ward et al., 2002	ePTFE and PVA	ePTFE: 1 μm PVA: 60 μm	Membrane implanted in rat subcutis	7 weeks
Marshall et al., 2004	HEMA hydrogels	35 μm	Hydrogel implanted in mouse subcutis	4 weeks
Updike et al., 2000	ePTFE	5-10 μm (Shults et al., 2006)	Glucose sensor implanted in dog subcutis	162 days ∗ (best of 6 sensors)
Yu et al., 2006	Epoxy-enhanced polyurethane	Not specified	Glucose sensor implanted in rat subcutis	56 days ∗ (best of 9 sensors)
Gilligan et al., 2004	ePTFE	Not specified	Glucose sensor implanted in human subcutis	185 days ∗ (best of 5 sensors)

NOTE: PTFE (polytetrafluoroethylene), ePTFE (expanded polytetrafluoroethylene), PVA (polyvinyl alcohol), HEMA (hydroxyethyl methacrylate).

∗

Maximum time sensor remained functional in vivo.

While porous biomaterials seemingly create the ideal environment for an indwelling glucose sensor, porous coatings applied to sensors have had less than ideal results. A critical factor in sensor failure in vivo is the fibrotic capsule that forms around glucose sensors (Dungel et al., 2008). Despite porous coatings stimulating the formation of vascular networks around glucose sensors in rats, newly formed vessels within porous coatings have been unable to overcome the diffusion barriers imparted by the collagen capsule (Dungel et al., 2008). Failing sensor sensitivity was found to correlate with increasing collagen deposition within the sponge implant (Dungel et al., 2008). Additionally, other factors, such as mechanical stresses imposed by the percutaneously implanted sensor, may have overshadowed the angiogenic-inducing, collagen-reducing properties of porous coatings (Koschwanez et al., 2008).

Recently, human studies (Gilligan et al., 2004) were performed using sensors covered with a porous angiogenic and bioprotective ePTFE membrane (Updike et al., 2000; Shults et al., 2006). Unfortunately, inflammation within the angiogenic layer in 80% of sensors, in addition to packaging failure in 60% of sensors, resulted in only 20% sensor survival after six months (Gilligan et al., 2004).

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Wireless biosensors for POC medical applications

M.S. Arefin , ... M.R. Yuce , in Medical Biosensors for Point of Care (POC) Applications, 2017

7.4.1 Wireless implantable glucose biosensors

Implantable glucose sensors with interface circuits hold a great potential for the continuous measurement and monitoring of blood glucose in patients with diabetes. The sensor can be implanted under the skin [4,33,34,43,112].

The functional block diagram of an implantable microsystem for blood glucose monitoring designed by Ahmadi and Jullien is shown in Fig. 7.10 [34]. The glucose sensor is an amperometric electrochemical biosensor generating a current from the electrochemical reaction between glucose and a glucose oxidase layer on working electrode (WE). The use of iridium-oxide nanoparticles helps for the transfer of the electrons from the glucose oxidase to WE. The reference electrode (RE) eliminates the potential arising from the solution medium. The counter electrode (CE) acts as a reference half-cell to supply the required current for the electrochemical reaction, whereas the WE act as a sensing half-cell to produce the current. The external reader inductively transfers power to the implantable microsystem and receives the transmitted measurement data of blood glucose concentration from the microsystem. The data transmission is performed for every 10   min using a load-shift keying modulation scheme. The interface circuit of the microsystem consists of an RF front-end circuit for receiving RF signals, rectifying, and generating the supply voltage, and a data acquisition circuit for converting the current from glucose sensor to pulse.

Figure 7.10. Block diagram of the implantable microsystem for continuous glucose monitoring.

M.M. Ahmadi, G.A. Jullien, A wireless-implantable microsystem for continuous blood glucose monitoring, IEEE Transactions on Biomedical Circuits and Systems 3 (2009) 169–180.

The cross-sectional view of the glucose biosensor is illustrated in Fig. 7.11. The titanium–nickel–gold–titanium metallization is essential for the WE, CE, interconnect traces, and bonding pads. The glucose oxidase on gold acts as a biologically sensitive layer. The silver metal layer at RE acts as an Ag/AgCl electrode, which generates current from the solution medium. The integrated interface circuit and the wireless transmitter are bonded on this wafer. The off-chip components and inductive coil for energy transmission are connected on this wafer. The dimension of the microsystem is 8   mm   ×   4   mm and its thickness is 1   mm.

Figure 7.11. Cross-sectional view of the implantable microsystem and glucose sensor for continuous blood glucose monitoring.

M.M. Ahmadi, G.A. Jullien, A wireless-implantable microsystem for continuous blood glucose monitoring, IEEE Transactions on Biomedical Circuits and Systems 3 (2009) 169–180.

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Glucose Biosensors

Dr Dennis Fitzpatrick PhD CEng BEng(Hons) , in Implantable Electronic Medical Devices, 2015

4.5 Implantable Glucose Sensor by GlySens

The implantable glucose sensor from GlySens is a long-term implant lasting for 1 year or more and does not require continuous calibrations. The implant consists of an integrated glucose sensor with signal conditioning circuits, a wireless telemetry circuit, and a 1-year lifetime battery, all housed in a hermetically sealed titanium housing ( Figure 4.6). The wireless radio frequency (RF) link communicates with an external receiver providing continuous glucose monitoring.

Figure 4.6. Implantable GlySens glucose sensor. Cross-sectional view shows electronics module (A), telemetry transmission portal (B), battery (C), and sensor array (D) (Gough et al., 2010).

(Reprinted with permission.)

The glucose sensor is an amperometric glucose sensor based on the detection of oxygen. The oxygen sensor incorporates dual-enzyme electrode technology with both enzymes, glucose oxidase and catalase, immobilized in a cross-linked protein gel. The catalase enzyme reduces the deactivation of the glucose oxidase enzyme in the presence of hydrogen peroxide, increasing sensor stability and effective lifetime. The reference oxygen sensor contains no enzymes.

An integrated three electrode potentiostatic circuit is used to set up the working electrode voltages and to measure the differential currents between the two oxygen sensors. The sensor array consists of four working-counter platinum electrode pairs and an Ag/AgCl reference electrode, microfabricated onto the surface of an aluminum disk measuring 12   mm in diameter. The enzymes are immobilized by cross-linking with albumin using glutaraldehyde into a gel and are covered with a protective semipermeable membrane layer of polydimethylsiloxane, reducing interference from unwanted molecules.

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Colloids for Nanobiotechnology

Jingyue Xu , ... Niko Hildebrandt , in Frontiers of Nanoscience, 2020

3.2.3 QDs as FRET donors and acceptors

QDs could engage in FRET with each other as a QD-to-QD FRET system. Compared with the common QD-to-dye and QD-to-fluorescent protein systems, a primary advantage of QDs in QD-to-QD FRET sensing is their extreme photostability compared to organic dyes or fluorescent proteins, making QDs uniquely suited for longitudinal studies, where measurements or images are taken repeatedly over extended periods of time. Their stability makes them good candidates for device-on-a-chip applications and for sensors designed for use outside of the laboratory setting. In addition, the extraordinary brightness of QDs, primarily due to their large absorption cross-section, yields considerable fluorescence output with relatively fewer emitters and potentially decreases the limit of detection in sensing applications. However, QD-to-QD FRET is challenging due to the broad, overlapping excitation spectra from the two nanocrystals, precluding selective excitation of the donor. This introduces cross talk and artificially creates a large background signal, which is a major challenge in QD-to-QD FRET sensor design. Due to their prominent color and intensity changes, QD-to-QD FRET signals from heterogeneous QD samples (two different-sized, but physically comingled QDs that almost always have the same composition) have been mostly adopted in sensing applications [94].

3.2.3.1 Small molecule sensing

A glucose sensor was designed using green CdTe QD-Concanavalin A (Con A) conjugates and red QD-NH ₂-glucosamine hydrochloride conjugates as a FRET pair [95]. Con A is a lectin with high affinity to manno- and gluco-oligosaccharides, including dextran. Prefilling of the binding sites on the CdTe QD-Con A conjugate by glucose inhibited FRET between the two QDs, allowing for indirect sensing. This FRET-based inhibition assay provided a fluorometric quantification method for glucose.

3.2.3.2 Antibody/antigen sensing

Despite the large sizes of QD-antibody conjugates, two different QDs could be conjugated to an antibody and antigen and then used within a FRET-based immunoassay format. Liu et al. used the specific binding of IgG as a secondary antibody to induce an interaction between donor and acceptor QDs [96]. Antihuman CD71 monoclonal antibody (anti-CD71) was conjugated with red QDs and was used to label HeLa cells successfully. Then green QD-labeled IgG was added to bind the red-QD-conjugated anti-CD71 on the cell surface by immunoreactions. This study not only proves that it is possible to use QD-to-QD FRET for studying interactions on living cell membranes, but also shows their potential valuable application to screening of antibody/antigen with bioactivity detected and identification of tumor.

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Application of responsive polymers in implantable medical devices and biosensors

J. Li , ... Z. Zhang , in Switchable and Responsive Surfaces and Materials for Biomedical Applications, 2015

11.12.1 Enzyme electrode-based biosensors

Most commercialized implantable glucose sensors are based on a glucose–enzyme reaction with their products detected by an amperometric sensor. The enzyme, usually a GOX, is immobilized on an electrode to provide a redox reaction with glucose and generate a current at the electrodes. With the immobilization of the GOX, the redox polymer layer can be viewed as a glucose-responsive system ( Heo & Takeuchi, 2013; Wang, 2007). Commercialized implanted glucose sensors are usually implanted subcutaneously with a sensor life span of less than 7   days. In addition to subcutaneous implantation, glucose sensors integrated with intravascular catheters (Armour, Lucisano, McKean, & Gough, 1990) and contact lenses (Yao, Shum, Cowan, Lähdesmäki, & Parviz, 2011) have also been reported. Due to the discomfort, short life span, and the requirement for daily calibration of current implantable glucose sensors, there is still a tremendous need for improving blood glucose monitoring (Gerritsen, Jansen, & Lutterman, 1999).

The polymers that immobilize enzymes on the electrodes respond to analytes through an electron-transfer process. Different materials and their combinations have been applied on the electrodes to improve the sensitivity, biocompatibility, and lifetime of the biosensors. The redox enzymes can be either chemically bound within the cross-linked polymer network or physically embedded within the materials. Figure 11.5 shows a typical implantable glucose electrode with a sandwich structure designed for immobilizing the enzyme (Clark & Duggan, 1982). Glucose diffuses through an outer layer to reach the immobilized enzyme, GOX, which is placed very close to the surface of a platinum electrode. The outer layer should have a function of allowing maximum passage of oxygen and retarding the passage of glucose to control glucose diffusion. This membrane must also be biocompatible and stable in vivo. The inner layer serves as the support for enzyme immobilization and also as the selective membrane for H₂O₂. The enzyme layer is where glucose is converted to the electoactive species H₂O₂.

Figure 11.5. A typical implantable glucose electrode that is responsive to glucose concentration.

Dialysis membranes from cellulose acetate (CA) are among the earliest substrates that have been used by Clark and Lyons (Clark & Lyons, 1962). To exclude interfering anions, a negatively charged perfluorinated ionomer Nafion™ membrane was alternatively deposited with CA on the electrode (Zhang et al., 1994). As an outer layer, Nafion™ also provides protection and improves biocompatibility. Several needle-type glucose sensors with Nafion™ layer remain functional for at least 10   days after subcutaneous implantation in dogs, without degradation of their sensitivity (Moussy, Harrison, O'Brien, & Rajotte, 1993; Moussy, Jakeway, Harrison, & Rajotte, 1994). Other examples of interference-excluding membranes include polydimethylsiloxane (PDMS) (Ertefai & Gough, 1989; Yang, Atanasov, & Wilkins, 1995), polyurethane (Moatti-Sirat et al., 1992; Pickup, Claremont, & Shaw, 1993; Yu, Long, Moussy, & Moussy, 2006), and modified CA (Qiu & Hu, 2013). These polymers allow for permeation of molecules having a similar molecular weight to the analyte; interferences of sizes larger than the analyte are excluded. Durability of the biosensors can be improved using these protective membranes. For example, an epoxy-enhanced polyurethane membrane was used as the outer protective membrane of the sensor and kept functioning in rats for 10–56   days (Yu et al., 2006). A glucose sensor, covered by a thin electrolyte layer, a protective layer of medical-grade PDMS, and a membrane of PDMS with wells for the immobilized enzymes located over certain electrodes was capable of performing long-term monitoring of tissue glucose concentrations by wireless telemetry (Gough, Kumosa, Routh, Lin, & Lucisano, 2010). The implanted sensor was functional in a pig model for more than 1 year, indicating significant progress in extending the life span of implantable glucose sensors (Gough et al., 2010).

An earlier study used CA as an inner layer to exclude molecules such as ascorbate, urate, and bilirubin (Clark & Duggan, 1982). The interference by small, electroactive compounds could be further reduced by incorporating conductive polymers. Conductive polymers have been coated on the implantable glucose sensor electrodes to provide efficient transfer of electric charge produced by the biochemical reaction to an electronic circuit. The conductive polymers, usually intrinsically conducting polymers with conjugated backbones, provide high electron affinity and are highly susceptible to chemical or electrochemical oxidation or reduction (Singh, 2012). Using enzymes during electrochemical polymerization, enzymes can be immobilized on the electrodes with conductive polymers. For example, enzymes entrapped within films such as polypyrrole (PPy), polyaniline, or polythiophene, prepared by electropolymerization from aqueous solutions, have been commonly used to prepare glucose electrodes. Like Nafion™, some conductive polymers can exclude interference of molecules with sizes larger than the analyte, such as overoxidized PPy (Rizzi, Centonze, & Zambonin, 2000), poly(o-phenylenediamine) (Malitesta, Palmisano, Torsi, & Zambonin, 1990; Sasso, Pierce, Walla, & Yacynych, 1990), and poly(quinone) (Arai, Shoji, & Yasumori, 2006; Kaku, Okamoto, Charles, Holness, & Karan, 1995). With some in vivo results for more than 10   days (Moussy et al., 1993, 1994), conductive polymers have shown activity, sensitivity, and selectivity, and possessed good durability on glucose electrodes.

Within the enzyme layer, GOX has been immobilized through cross-linked albumin, synthetic hydrogels (Guiseppi-Elie, 2010), and conductive polymers (Singh, 2012). Hydrogels are commonly used materials applied to the electrodes to immobilize enzymes and/or provide biocompatibility, permeability, and fouling resistance. PolyHEMA and its copolymers are among the earliest hydrogels to be applied on electrodes (Shaw, Claremont, & Pickup, 1991). To improve porosity, hydrophilicity, or biocompatibility, 3-dihydroxypropyl methacrylate (DHPMA) (Wang et al., 2008; Yu et al., 2008), N-vinyl pyrrolidone (Heineman, 1993), vinyl alcohol (Vaddiraju, Singh, Burgess, Jain, & Papadimitrakopoulos, 2009), ethylene glycol (Quinn, Pathak, Heller, & Hubbell, 1995), 2-methacryloyloxyethyl phosphorylcholine (MPC) (Chen et al., 1992), and carboxybetaine (Yang, Xue, Carr, Wang, & Jiang, 2011; Zhang et al., 2009) have all been applied. These hydrogels are ionically but not electronically conductive and usually demonstrate high interfacial impedances. Conductive polymers are incorporated into the hydrogel network to improve the stimuli responsiveness and reduce interfacial electrical impedances (Heller, 2006) (Guiseppi-Elie, 2010). Incorporating an osmium complex within the hydrogel network could also improve the redox efficacy (Kenausis, Taylor, Katakis, & Heller, 1996; Mano, Mao, & Heller, 2005).

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Biosensors Development

Azrul Azlan Hamzah , Sh. Nadzirah , in Reference Module in Biomedical Sciences, 2022

Healthcare biomedical application

Over decades, glucose sensors and pregnancy tests are the most frequently reported and successful commercial biosensors. The most impressive development of a glucose meter is the non-invasive glucose sensors instead of finger pricks. They have been approved by the FDA and commercialized such as FreeStyle Libre (wear at back arm), Eversense CGM (implant), Dexcom G6 CGM (wear at abdominal area), and Guardian Connect System (abdominal area) ( Cherney, 2021). These sensors are able to be connected with smartphones. Further improvement may focus on improving the sensitivity of the non-invasive glucose meter, increase options for remote monitoring through smartphones and smartwatches, and direct measurement of blood sugar through smartwatch.

Oximeter is a non-invasive optoelectronic biotransducer that measures the oxygen level (deoxygenated and oxygenated) hemoglobin in red blood cells (RBC). Currently, it is designed in a tiny shape that is usually clipped over a finger and the red and infrared light emitting diode of the device will pass through the tissue. The oximeter then reads the refraction of the light and converts it into pulse wave on the monitor (Organization, 2011). The saturation of oxygen in hemoglobin is prescribed by the SpO₂ value. For SpO₂ value below 95%, it is advisable to consult with medical professionals. This sensor then evolves as a wearable sensor which has been installed in smartwatches like Garmin, Fitbit, and various smartphones worldwide (Kooistra, 2018; Browne et al., 2021). This home-healthcare benefits for monitoring of overdoing exercise, also for individuals that are affected by respiratory issues during the COVID pandemic.

Due to the current COVID-19 pandemic, a direct and simple handling biosensor is urgently required. Current COVID tests are based on antigen analyte or RNA analyte. Antigen assay COVID test is much faster and cheaper than the gold standard of Real-time polymerase chain reaction (RT-PCR) (Guglielmi, 2020). However, WHO has not recommended the use of the antigen-detection approach for COVID-19 as it has sensitivity issues (WHO, 2020). Yet, they still encourage scientists to study the performance and potential diagnostic utilities. PCR tests are almost 100% accurate in spotting COVID-19 viruses. However, it is expensive due to the high-end equipment involved and must be handled by well-trained personnel. There are more than 350 RT-PCR COVID-19 test kits available until recently, and some of them have been approved by the United States Food and Drug Administration (US-FDA) (Garg et al., 2021). In addition, kits with CE-IVD certification attract more manufacturers and distributors. Until recently, the average sensitivity of kits is 95% LOD which are based on C_t value. C_t stands for the cycle threshold; number of cycles required to amplify the viral RNA of COVID to a detectable level. Test with C_t  ≤   34.5% is considered as positive. Table 1 shows the state-of-the- art of biosensors that has been commercialized and under research.

Table 1. Shows example of state-of-the- art human direct-interfaced biosensors for various target detection.

No.	Sensor	Product type	Target	Performance
1.	FreeStyle Libre	Arm-based	Glucose	Accuracy: 9.3% MARD
2.	Eversense CGM	Implant	Glucose	Accuracy: 8.5–9.6% MARD
3.	Dexcom G6 CGM	Abdominal-based	Glucose	Accuracy: 9% MARD
4.	Guardian Connect System	Wear at abdominal	Glucose	Accuracy: 8.7% MARD
	GraphWear	Torso-based	Glucose	NA
5.	Kenzen Wear's Echo patch	• Torso-based	Sodium, potassium, glucose	NA
	Salixium-COVID-19 Rapid Antigen Rapid Test	• Saliva swab or nasal swab	COVID-19 antigen	Sensitivity >   91%
6.	LYHER® antigen test kit (Colloidal Gold)	Oral fluid	N protein of SARS-CoV-2	Sensitivity 93.3%
7.	Halo H1	Wrist-based	Fluid level (hydration)	NA
8.	MightySat™ Rx (Masimo)	Finger-based	Oxygen saturation (SpO2)	Sensitivity <   90% Specificity >   90%
9.	Pulsox-310 (Konica Minolta)	Finger-based	Oxygen saturation (SpO2)	Accuracy: 70–100%

Researchers are urgently studying the development of electrical and optical biosensors (Chaibun et al., 2021). The end stage should be wearable, hence easy to personally or remotely monitor health conditions, fast tracking of disease symptom and health recovery of an individual. Current wearable sensors that are widely used such as smartwatches are able to monitor resting heart rate and fever (Ke et al., 2019; Ranjan et al., 2019; Purohit et al., 2020a).

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Strategies to improve the hemocompatibility of biodegradable biomaterials

P. Mulinti , ... A.E. Brooks , in Hemocompatibility of Biomaterials for Clinical Applications, 2018

10.4.3 Polymer hybrids

Implantable sensors, including glucose sensors, play a significant role in current medical practice; however, their indwelling lifetime is often limited by protein adsorption, which progresses to a chronic FBR. Although several groups have tried coating these sensors with a variety of synthetic polymers such as PU, poly(2-methoxyethyl acrylate), and PVA [ 110], their durability is still limited. A material that is compatible with blood while still allowing signal molecule (e.g., glucose) transport is needed. Recently, a polyester fabric was coated with a biodegradable PGA sheet and implanted in a rat model. As the PGA degraded, cells deposited their extracellular matrix, much of which is collagen, on the polyester fabric. This hybrid fabric was decellularized, and its mechanical compliance and hemocompatibility were evaluated, revealing a material that resisted platelet adhesion and thrombus formation [111].

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Molecular Events at Tissue–Biomaterial Interface

Themis R. Kyriakides , in Host Response to Biomaterials, 2015

Molecular strategies to enhance angiogenic responses

Specific applications, such as glucose sensors and tissue engineered constructs, have increased dependence on angiogenesis. Interestingly, prolonged inflammation is often associated with persistent angiogenic responses but this is not the case with implanted biomaterials. Conceivably, this relative decrease in angiogenesis is due to loss of production of pro-angiogenic factors by late stage macrophages or sequestration of these factors outside of the avascular fibrous capsule, where an increased angiogenic response is often seen. Alternatively, excessive deposition of angiogenesis inhibitors during matrix production could also negatively influence angiogenesis. Regardless of the mechanism, the result is inefficient transport of molecules from microcirculation to the implant. Several groups have attempted to increase the number and stability of vessels in the FBR. Such strategies include delivery of pro-angiogenic factors such as VEGF, PDGF, and MCP-1, which constitute the majority of neovascularization approaches ( Jay et al., 2010; Richardson et al., 2001; Brudno et al., 2013; Klueh et al., 2005). Furthermore, the secretion and sequestration of angiogenic factors by the ECM can be replicated closely by modulating their release. This controlled release could be accomplished by engineering chemical or enzymatic susceptibilities that allow for spatial and temporal control of release. In addition, engineered enzyme (MMP)-sensitive hydrogel systems were shown to stimulate vascular formation (Seliktar et al., 2004; Kraehenbuehl et al., 2008). Alternatively, targeting the expression of anti-angiogenic factors, such as TSP2 or prolyl hydroxylase domain protein 2 (PHD2), was shown to enhance vessel density in the FBR (Kyriakides et al., 2001a; Nelson et al., 2014). Specifically, it was shown that gene-activated matrix delivery of an antisense TSP2 cDNA enhanced blood vessel formation and altered collagen fibrillogenesis in mouse subcutaneous implant models. Similarly, Figure 5.7 shows delivery of PHD2-specific siRNA from a porous polyester urethane (PEUR) scaffold that resulted in sustained increased blood vessel formation associated with increased VEGF and bFGF.

Figure 5.7. Sustained silencing of PHD2 increases angiogenesis within PEUR tissue scaffolds. CD31 staining was significantly increased within PHD2 scaffolds at day 14 and day 33 (scale=200   μm, vessels appear red, nuclei are counterstained purple with hematoxylin, and the white space represents residual PEUR scaffold). (F) Micro-CT images visually demonstrate the increased vasculature within the PHD2-NP scaffolds.

Reprinted with permission from Nelson et al. (2014).

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