Diabetes mellitus is a chronic disease, that occurs when the pancreas does not produce enough insulin, or when insulin is not effectively used. Insulin is a hormone that regulates blood sugar level. This leads to an imbalance in concentration of glucose within the blood; causing either hyperglycaemia or hypoglycaemia. This over a period of time leads to serious damage to the nerves and blood vessels. It has been noted to affect more than 346 million people worldwide (WHO, 2012) and spending on complications caused by diabetes, cost the NHS almost £ 10 billion in 2011(Diabetes UK, 2012). These complications can either be acute or chronic, both categories cause severe disadvantages to the patient. By 2035 it is predicted that the NHS will spend £16.8 billion, 17% of its entire budget for diabetes treatment. (NHS, 2012)
Figure 2.1 shows significantly rising incidence of diabetes in England (Diabetes UK, 2012)
Two general types of diabetes:
Type 1, Insulin-dependent diabetes mellitus (IDDM) generally afflicted at a younger age(between ages 10 and 16) (Harvard Health Publications, 2012) also known as juvenile diabetes mellitus and it is caused by the inability of the beta cells (in the pancreas) to produce or secrete insulin (figure 2). Approximately 10% of diabetics have type 1 (Newman and Turner, 2005).
Type 2, Non-insulin dependent diabetes mellitus (NIDDM) which makes up the majority of diabetes, 90% (Guyton and Hall, 2005) affliction usually occurs after the age of 40 years and it more prone to occur due to lifestyle. Obesity greatly increases the risk of diabetes (Harvard Health Publications, 2012). In this type, insulin has an inability to bind to cell receptors for the uptake of glucose (figure 3).
type 2 diabetes.jpg
type 1 diabetes.jpg
Figure 2.2 (left) and 2.3 (right) shows the physiological difference between type 1 and type 2 diabetes, respectively (Harvard Health Publications, 2012).
A biosensor is a compact analytical device or unit which transduces biological or biologically derived stimuli, be it physical or chemical, into readable signal for monitoring (figure 4). (Chambers et al., 2008) There are three main parts of a biosensor:
The biological recognition elements that distinguish the target molecules in the presence of various chemicals
A transducer that converts the biorecognition event into measurable signal
A signal processing system that converts signal into a readable form
Figure 2.4 shows schematic presentation of a biosensor (Belluzo et al., 2008)
There is a plethora of different biosensors developed, which use different sensing modalities like electrochemical, optical, thermometric, piezoelectric and magnetic (Yoo and Lee, 2010). The most researched being glucose biosensors which predominantly use an electrochemical sensing modality, characteristically implied in having simple operation, suitable sensitivity, reliable, quick and low manufacturing cost (Newman and Turner, 2005).
2.2.1 Glucose Biosensors
Glucose is the primary fuel for cells of most living organisms and commercially, it is used as a precursor for the production of molecules such as vitamin C, citric acid, gluconic acid, polylactic acid, sorbitol and bio-ethanol. Such large scale utilization of glucose necessitates the call for efficient feedback control systems, for which the use of glucose sensors is mandatory.
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Glucose biosensors, as mentioned earlier, predominantly use an electrochemical sensing modality. Electrochemical sensors measure the electrons produced as a by-product during an electrochemical reaction. This is done by targeting a main characteristic of most enzymes, whereby a redox reaction is performed, when catalyzing a substrate (Ratner et al., 2004; Heller and Feldman, 2008). Electrochemical sensors may be subdivided into Amperometric, Potentiometric, Conductometric and Ion-selective field-effect transistor (ISFET) based sensors (Li et al., 2007).
Figure 2.5 shows a chronological history of glucose biosensors. (Yoo and Lee, 2010)
Enzymatic Amperometric glucose biosensors are the most common devices commercially available, and have been widely researched over the last few decades (Yoo and Lee, 2010). Amperometric sensors comprise of two to three fundamental electrodes; a working electrode, a reference electrode and a counter electrode. The enzyme involved in catalyzing the target reaction is immobilized on the working electrode surface and the result of the biochemical reaction can then be found at the anode or cathode.
Figure 2.6 shows the components of a coil-type electrochemical, Amperometric glucose biosensor (Yu et al., 2005).
The enzyme used for the above glucose biosensor assembly is Glucose Oxidase (GOx). This is one of two enzyme families that are used as the recognition element for the electro-oxidation of glucose, the other family being PQQ-glucose dehydrogenase (PQQ-GDH).
Although both enzymes can be used for electrochemical glucose biosensor, GOx has been more widely used for in vivo glucose biosensor research. This is due to GOx being more specific to glucose compared to PQQ-GDH. In the electrochemically relevant half-reaction, glucose is oxidized by GOx at the rate of 5×103 glucose molecules per second whereas PQQ-GDH oxidizes glucose at rate of 11,800 glucose molecules per second (Heller and Feldman, 2008).
2.3 The importance of blood glucose Monitoring
Blood glucose measurement has been highlighted for a diabetes treatment; it has been shown that effective monitoring of blood glucose can lead to 30-70% decrease in progression of the disease. (Diabetes control and complications trial research group, 1993) With careful administration, these complications can be delayed and even prevented. (Newman and Turner, 2005) Thus the American Diabetes Association recommends Self-Monitoring of Blood Glucose (SMBG) for Type 1 diabetics at least four times a day (American Diabetes Association, 1994 cited in Newman and Turner, 2005) while Type 2 diabetics should monitor their glucose levels twice a day. Blood glucose monitoring are undertaken using commercial ex-vivo/in-vitro glucose biosensors.
The commercial glucose biosensors available for home and clinical use are generally classified into two categories, self-monitoring of blood glucose (SMBG) and continuous glucose monitoring (CGM) devices. The former requires finger to be pricked to draw blood which is used on test strips to detect blood glucose concentration while the later involves non-invasive, minimally invasive or invasive methods for CGM. However, commercially only invasive devices like needle type electrochemical glucose biosensor are available which is inserted in the dermal or subcutaneous skin space and function for a maximum period of 7 days.
Figure 2.7 shows the advantage of CGM over SMBG (Medtronic, 2012)
The figures even with compliance of testing frequently are a snapshot for the fluctuating blood glucose measurements present, thus missing episodes of hypo or hyper-glycemia. Also various other aspects need to be taken into consideration such as what and when the patient last ate; patients exercise regime and the amount of medication that has been taken, cost of glucose test strips. Hence due to the various complexities, it appears that the frequency of SMBG testing is not at the required level in many parts of the world. (Newman and Turner, 2005) There is demand for improved designs of glucose monitors for better diabetes treatment which all glucose biosensor company’s research and design. Focus are highly involved in developing glucose biosensors for a long term Continuous Glucose Monitoring (CGM) system which will provide better understanding and monitoring of the fluctuating glucose levels for better individual treatment. It is also the bottleneck for the design of an artificial pancreas.
2.4 Continuous Glucose Monitoring
Continuous Glucose Monitors (CGM) measures interstitial glucose levels continuously and updates the glucose level every 1 to 5 minutes. The CGM system consists of the following:
A monitor to display the information
Sensor that is usually inserted in the subcutaneous tissue using sensor inserter
Figure 2.8 shows Continuous Glucose Monitoring System where A -Sensor, B- Sensor inserter, C- Monitor, D- Monitor connected to a docking station to download data into a computer (Gross et al., 2000).
A transmitter that transmits the sensor data to the monitor.
CGM can provide both retrospective as well as real-time information for the detection of hypoglycemic and hyperglycemic excursions, prediction of impending hypoglycemia and wide fluctuations in glucose levels also known as glycemic variability (Yee and Klonoff, 2010).
Therefore they are beneficial to both the physician and the patient for trend analysis, better insulin treatment and for long term reduction to metabolic complications. However the major goal of CGM devices still focuses upon its impact within the diabetic society and for the commercial production of an artificial pancreas, which monitors and delivers effective therapy to fluctuations in blood glucose to maintain the healthy range within a diabetic, preventing acute and chronic complications associated with erratic blood glucose levels within patients.
The CGM systems available commercially include Guardian REAL-Time, Minimed Paradigm REAL-Time, Minimed CGMS system Gold, SEVEN by Dexcom (San Diego, CA, USA) and Freestyle Navigator by Abbott (Abbott Park, IL, USA). All these systems are based on enzymatic, amperometric glucose biosensors which measure glucose in the interstitial fluid (ISF). While the Glucoday system is based on the principle of microdialysis, where ISF is drawn using minimally invasive dialysis tubing to an amperometric device located outside the body. This device is reasonably bulky and is often limited to hospital. The reliable use of these devices for CGM is limited to 2 to 7 days, due to concern related with biocompatibility.
Therefore there is a need for better CGM technology, with aspiration for producing devices with a longer life span. The requirement for improving longevity is initially for improved glucose monitoring, but also for implicating the potential of designing an effective artificial pancreas, with a closed loop system. (Hovorka, 2008)
Glucose biosensors have been shown to fail for various reasons, but one large area of performance failure in vivo, is the problem of sensor biocompatibility. Previous researchers have focused on altering the outer membrane component of glucose biosensors for improving biocompatibility, as the outer membrane acts as the interface between the functional electrochemical components of the glucose biosensor and the external biological environment. Therefore developing an outer membrane with better surface properties will control and prevent negative biological responses. (Wisniewski et al, 2000)
2.5 Biocompatibility – the bottleneck for implantable glucose biosensors
Wound healing response
Figure 2.9 shows host response to the insertion of a biomaterial (Chan et al., 2008)
Secretion of ECM by fibroblasts and capillary formation
Macrophages and FBGCs
Monocytes differentiate to macrophages
Host tissue in-growth, Implant degradation and resorption
Macrophages fusion to form FBGCs, Frustrated phagocytosis
Proliferation of fibroblasts and endothelial cells
Fibrous capsule formation
Resolution of injury
Implant site remodelling
Host tissue replacement/regeneration
Protein adsorption, Water adsorption, Implant degradation
The outer membrane component of glucose biosensors is generally composed of polymer material, most frequently a synthetic polymer, known to have better durability, mechanical strength, viscoelastic properties and ease for manufacture and processing. However they lack biocompatibility when compared to natural polymers. (Wang et al, 2004; Liang et al, 2007)
Therefore through the use of electrospinning, a process capable of fabricating nanofiber mesh-like membranes advantageous for improved outer membrane characteristics as well as bio-mimicry of biological components, such as the extracellular matrix (ECM). (Sill and Von Recum, 2008)
Electrospinning is a non-woven fibre spinning technology that allows spinning of fibres having diameter ranging from 2 nm to 10s of µm (Bhardwaj and Kundu, 2010). Creating fibre meshes, in the nanometre and micrometre range, is important for biological application, as small pore sizes created in nanofibres compositions, prevent the adherence of certain cells. (Sill and Von Recum, 2008) Therefore the production of nanofibres meshes would be advantageous for the outer membrane of in vivo glucose biosensors for longevity of the sensor in vivo.
The electrospinning process consists of a polymer solution held by its surface tension at the end of a capillary tube which is subjected to an electric field. Here the
charge is induced onto the liquid surface by an electric field. As a result, mutual charge repulsion causes a force which is directly opposite to the surface tension. And as the intensity of the electric field is increased, the hemispherical surface of the solution at the tip of the capillary tube elongates to form a conical shape known as the Taylor cone. When the electric field reaches a critical value at which the repulsive electric force overcomes the surface tension force, a charged jet of the solution is ejected from the tip of the Taylor cone. The trajectory of the jet can be controlled by an electric field as it is charged. When the jet travels in air, the solvent evaporates, leaving behind a charged polymer fibre which is laid randomly on a collecting surface. Thus continuous fibres are laid to form a non-woven fabric. (Doshi and Reneker, 1995)
Figure 2.9 Schematic diagram shows the general apparatus used for electrospinning. (Ziabari et al, 2009)
Figure 2.10 Photographs illustrating the instability region of a liquid jet electrospun from an aqueous solution of poly(ethylene oxide) PEO A) conventional and B) high speed camera with exposure times of 1/250 s and 18 ns respectively. (Li and Xia, 2004)
In real-time, the bending instability of the jet appears to naked eye or using conventional photography as the single stream of polymer solution splitting into large number of jets in a cone shape. However, using high speed camera, Shin et al. demonstrated that it is basically a single fibre that bends and whips rapidly to give the illusion of splitting and multiple jets (Shin et al., 2001)
2.5.1 Electrospinning apparatus
The apparatus required for electrospinning is simple in assembly and consists of a high voltage power supply, a syringe pump with tubing to transport solution from the syringe to the spinneret, and a conducting collector. The spinneret and collector are aligned opposite to each other and together they are usually oriented either horizontally or vertically. Further variations in the equipment can be achieved with various spinneret and collector combinations.
Spinneret constructions are used to achieve greater versatility in the nano-fibre structures. They can be categorised into single or multiple and needle-type or needleless configurations (Teo and Ramakrishna, 2006). The usual electrospinning setup consists of a single blunt end needle for spinning single polymer fibres (figure 2.11a). To obtain composite fibre structure, either coaxial (Figure 2.11b) or dual-capillary (side-by-side) (Error: Reference source not found 2.11e) spinnerets are used. The former produces a core-shell while the latter produces side-by-side fused fibre structures respectively. The advantage of using coaxial spinnerets is to obtain fibres from incompatible or non-electrospinnable polymer solutions (Teo and Ramakrishna, 2006). For instance, a non-electrospinnable polymer solution can be extruded in the inner capillary, while the spinnable solution extruded in the outer capillary. Thus the inner solution is enclosed by the outer solution resulting in a core fibre. Next the outer solution can then be dissolved to obtain the fibres of the non-electrospinnable solution. This method can also be used to produce hollow fibres by dissolving the inner core fibre. Overall the coaxial fibre results in a reinforced core polymer fibre encapsulated by a outer covering of a different polymer, whereas in the side-by-side fused composite fibres there are two different polymers placed alongside. Using various permutations and combinations of two polymers different membranes can be obtained for varied applications, particularly in the promising field of tissue engineering. In addition, the coaxial spinnerets can also be further modified to incorporate, for example, three concentric polymer fibres (Figure 2.11c), or a large fibre reinforced with three different polymer fibres within its composite structure (Figure 2.11d).
Figure 2.11 shows needle-based spinnerets and their cross-sections, a) single needle, b) two concentric needles, c) three concentric needle, d) three side-by-side needle inside a large needle and e) side-by-side needles. The numbers 1 to 4 indicate different polymer solutions (Teo and Ramakrishna, 2006).
In fundamental configuration, the spinneret needle is aligned perpendicular to a static grounded flat plate collector. During electrospinning process, the seamless fibre gets collected as a sheet, usually on an aluminium foil placed on top of the flat plate collector. The resultant sheet has random fibre orientation. However, the orientation, 3D architecture and properties of the electrospun structures can be varied by the application of an electric field between the spinneret and collector or by rotating the collector. In addition, the use of alternating-current (AC) high voltage supply instead of the traditional direct-current (DC) high voltage supply for charging the electrospinning solution is reported to induce better alignment of the fibres. (Kessick et al., 2004)
Various collector configurations used in research are illustrated below in figure 2.12. (Teo and Ramakrishna, 2006)
Large area of aligned fibres
Highly aligned fibrous assemblies difficult to fabricate
If rotating speed is too high fibre breakage may occur
Easy to obtain highly aligned fibres
Easy transfer of aligned fibre to another subtrate
Thick layer of aligned fibre are not possible
Limit in length og aligned fibres
Rotating wire drum collector
Highly aligned fibres are achievable
Cannot achieve thicker layer of aligned fibre
Fibres may not be aligned right through the whole assembly
Drum collector with wire wound on it
Highly aligned fibres possible
Area of aligned fibres on the wire can be adjusted by varying wire thickness
Aligned fibres are concentrated on the wire instead of the whole drum
Rotating tube collector with knife-edge electrodes below
Highly aligned fibres
Whole tube covered by aligned fibres
Thick layer of aligned fibre deposition possible
Set-up requires negative electrode to be effective
Only small diameter tube possible
Highly aligned fibres achievable
Direction of fibre alignment on tube controlled
Thicker layer of aligned fibre deposition
Set-up requires negative electrode to be effective
Only small diameter tube possible
Highly aligned fibres achievable
Fabrication of arrayed fibres by attaching rotatable table on edge of disc possible
Incapable to retain high fibre alignment at same rotating speed when deposited fibres are thicker
Small area of fibre alignment
Array of counter-electrodes
Inconsistent fibre pattern throughout assembly
Limited Assembly area
Thicker fibre assembly not possible
Rotating drum with sharp pin inside
Fabrication of large area of arrayed fibres
Thicker area of arrayed fibre assembly not possible
Blade placed in line
Highly aligned collected yarn
Limited fabricated yarn length
Deposited fibres need to be dipped in water before yarn formation
Ring collector placed in parallel
Fabrication of twisted yarn
Limited fabricated yarn length
One ring needs to be rotated to twist fibres that deposit into yarn
Controlled deposition using ring electrode
Area of fibre deposition can be minimized
Rings have to be given positive charge
Figure 2.12 Table showing a variety of collector configurations that have been tested for influencing fibre deposition on the collector. (Teo and Ramakrishna, 2006)
Fibre deposition can be controlled over an area
Rings have to be given positive charge
Area of fibre deposition is large although confined within the ring
2.5.2 Electrospinning Parameters
The many parameters which affect and/or control the process of electrospinning of fibres and their resulting morphology and diameter are as follows:
Polymer molecular weight
Solution surface tension
Distance of electrode source from the target substrate
Solution flow rate
All the variables mentioned above not all are neither fundamental control parameters nor they independent of each other. For example, solution viscosity is a function of both concentration and molecular weight. (Shenoy et al., 2005)
The electrospinnability of a polymer depends on its solution viscosity and electrical resistivity. If the viscosity is too low, the cohesiveness between the polymer chains (surface tension) is low to hold the fluid jet emerging from the Taylor cone together. Similarly, if the solution is too conductive, the repulsive forces between the polymer molecules cause the fluid jet to breakdown, forming droplets. With increasing viscosity and electrical resistivity, the breakup of the jet into droplets transitions into the formation of beads-on-string fibre structure before proper fibres form (Shenoy et al., 2005). Further increase in viscosity results in increasing fibre diameter until a maximum viscosity beyond which the Taylor cone becomes too big, causing the jet to become unstable (Megelski et al., 2002, Jarusuwannapoom et al., 2005, Demir et al., 2002, Deitzel et al., 2001a). Thus, for a given polymer, there exists a range of solution viscosity, wherein proper fibres form. This range varies for any specific polymer depending on the solvent it is dissolved in and also varies for different polymers (Huang et al., 2003). The solution parameters not only determine the electrospinnability of a polymer, but also affect the fibre morphology and diameter.
Materials and Methods
Electrodes: Platinum-iridium (10IR5t) (Pt:Ir, 9:1 weight ratio) and silver wires (AG5T), each having a diameter of 0.125 mm and covered with an insulating Teflon-coating, were obtained from ABC
Sensor Coatings: Bovine serum albumin (BSA), glutaraldehyde (GTA) grade I (50%), glucose oxidase (GOD) (EC 188.8.131.52, Type X-S, Aspergillus niger, 157,500U/g, Sigma), ATACS 5104/4013 epoxy adhesive, non-ionic surfactant Brij 30, polyurethane (PU) Z1A1, ,Tetrahydrofunan (THF) (THF, ACS reagent, >99.0%)
Sensor Function Testing: D-(+)-Glucose and phosphate buffered saline (PBS) tablets (0.01 M containing 0.0027 M potassium chloride and 0.137 M sodium chloride, pH 7.4, at 25 °C) were purchased from Sigma-Aldrich-Fluka, UK.
A miniature coil-type implantable glucose biosensor, developed in Moussy’s group (Yu et al., 2005), is used as model sensor in this study. The amperometric sensor is a two electrode system based on Pt/Ir working and silver/silver chloride (Ag/AgCl) reference electrodes.
Figure shows the components of implantable glucose biosensor
Working electrode: It was prepared by firstly removing 1 cm length of Teflon covering at either ends of the 8 cm long Pt-Ir wire having 0.125 mm diameter. On one end, the naked Pt-Ir wire was wound around 18 gauge needle (1/2 inch, BD) to make the working electrode coil. The core of the coil was then filled with cotton to enhance the immobilization of enzyme and also to prevent formation of an air bubble. The cotton reinforced coils were then cleaned using absolute ethanol, as well as by being placed in deionised water and sonicated using an ultrasonic bath (1510 Branson) for 30 minutes. The sensors were then heated at 60 °C for 20 minutes before immobilization of glucose oxidase (GOD) took place.
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Enzyme Loading: For immobilizing GOD on the working electrode coil, 1.0-1.5 µl of enzyme solution (39.3 mg/ml BSA, 8.2 mg/ml GOD and 1.6% (v/v) GTA (50%, v/v) dissolved in DI water) was loaded onto the cotton reinforced Pt-Ir coil and allowed to dry at room temperature for 30 minutes. This enzyme loading procedure was repeated 3 times for each working electrode. After the last coating, the enzyme layer was left overnight to allow GTA cross-linking to complete.
Mass-transport limiting membrane: To coat epoxy-PU (EPU) mass-transport limiting membrane, 1.5 µl of the EPU loading solution (26.7mg of PU, 8.9mg each of Part A and Part B of epoxy adhesive and 1 µl of Brij 30 dissolved in 4ml THF) was applied on the enzyme layer (Yu et al., 2007). The epoxy-PU coating has been shown to enhance function and longevity of the sensors (Yu et al, 2006). After air-drying at room temperature for 30 minutes, the solvent cast EPU layer was cured in an oven at 80oC for 20 minutes.
Reference electrode: The reference electrode was prepared by firstly stripping 1 cm of Teflon coating from both ends of a 7 cm long Teflon covered silver wire having 0.125 mm diameter. One end was carefully wound around a 30 gauge 1/2 inch hypodermic needle to get the coil end. The Ag coil was treated with ammonia solution for 30 seconds followed by 10 seconds in 6 M nitric acid. The coil was then washed in DI water and electroplated in 0.01 M HCl at a constant current of 0.1 mA using a galvanostat (263A, Princeton Applied Research, TN, US) for 5-6 hours in a 0.01M HCl solution of deionised water (100 ml) and Hydrochloric acid (100 μL). Electroplating consisted of the connection of the reference electrodes and a counter platinum mesh electrode within the beaker. The resulting Ag/AgCl reference electrode coils were rinsed with DI water.
Sensor Assembly: The sensors were assembled by inserting the Pt-Ir wire of working electrode through the coil of reference electrode until the two electrodes were separated by 5 mm and the Teflon covered part of the two wires were entangled along their entire length.
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