Ocean Optics32000-FLCCD Detector

Light source

Optical fiber

OZ OpticsBeam splitter

Sencil

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Abstract— Sencils™ (sensory cilia) are chemical sensors that are minimally invasive, disposable and easily readable to make frequent measurements of various analytes in vivo over a period of 1-3 months. A percutaneous optical fiber permits precise, reliable photonic measurement of chemical reactions in a nano-engineered polymer matrix attached to the internal end of the fiber. The first Sencils sense interstitial glucose based on measurement of fluorescence resonance energy transfer (FRET) between fluorophores bound to betacyclodextrin and Concanavalin A (Con A) in a polyethylene glycol (PEG) matrix. In vitro experiments demonstrate a rapid and precise relationship between the ratio of the two fluorescent emissions and concentration of glucose in saline over the physiological range of 0-500mg/dl. Chronic implantation in pigs has demonstrated biocompatibility. The Sencil platform can be adapted to detect other analytes in interstitial fluids. Keywords— Optical fiber, chemical sensor, fluorescence resonance energy transfer, affinity-binding assay, glucose, Quantum dots 1 Introduction

We are developing a platform technology that will support a family of disposable, minimally invasive, in vivo sensors that can measure various analytes in a patient over a period of 1-3 months. The key element is a chronically implanted optical fiber that provides a stable and reliable percutaneous port by which to interrogate the photochemical assay. A nanoengineered matrix is attached to the internal end of the optical fiber (Figure 1) that is implanted 1-2cm under skin to sample interstitial fluid. We call this a Sencil, for “sensory cilium”. The fist application of this technology is in glucose sensing.

In a previous paper1, we described the challenges of chemical sensing in vivo and the preliminary design and performance of a prototype Sencil. In this article, we describe important refinements to improve signal quality and mechanical stability.

Design and Fabrication of a Disposable, Percutaneous Glucose Sensor

Kuo-Chih Liao1,2, Thieo Hogen-Esch3,4, Frances J. Richmond1, Laura Marcu2, William Clifton1,2, Gerald E. Loeb1,2 1Alfred E. Mann Institute for Biomedical Engineering, 2Dept. of Biomedical Engineering,

3Dept. of Chemistry, 4Loker Hydrocarbon Research Institute, University of Southern California, CA, USA

Figure 1: Chemical sensing system. (A) Sensor components and relationships to tissue in vivo. (B) Photonic instrumentation for testing prototypes; will be miniaturized for portable clinical reader.

A B

Optical Fibers and Sensors for Medical Diagnostics and Treatment Applications VI, edited by Israel Gannot, Proc. of SPIEVol. 6083, 60830V, (2006) · 1605-7422/06/$15 · doi: 10.1117/12.643990

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Figure 2: Changes in FRET between fluorophores covalently immobilized on the flexible PEG matrix depend on changes in the distance between them, which in turn depends on the competitive natural affinity between Con A and various saccharides such as betacyclodextrin and glucose.

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2 Requirements and designs

2.1 Selectivity and sensitivity Analytes of clinical interest are mixed with a huge number of other biochemicals in the body whose

concentrations may vary widely and unpredictably. Chemical sensors (enzymatic or colorimetric method) generally have higher selectivity than optical sensors (absorption, polarimetry or Raman spectrum detection). Affinity binding between protein/receptor and ligand provide the desired specificity2; purely optical methods tend to be confused by biochemicals that share similar optical characteristics.3 The catalytic activity of enzymes such as glucose oxidase can provide a strong signal to detect but the reaction rates tend to drift independently of analyte concentration as the reactants deteriorate over time or there are changes in pH and temperature.4,5

The Sencil modifies the much-studied fluorescence resonance energy transfer assay (photochemical method) based on the recognition between jack bean lectin Concanavalin A (Con A) and saccharides.6 Succinylation of the Con A dimer7 stabilizes its affinity in bodily fluids; each monomer has a potential N-terminal binding site where a fluorophore can be attached and a nearby (2nm away) recognition site for saccharides such as glucose and betacyclodextrin .8 Betacyclodextrin binds to Con A, but can be displaced by glucose (figure 2). As described previously, we chose the relatively small and rigid betacyclodextrin because it had a lower affinity to Con A than other polysaccharides, permitting us to use a higher concentration in the matrix and still have proportional displacement over the physiological range of glucose concentrations1. Because the assay is a reversible reaction, it does not consume glucose as does glucose oxidase, which may produce a false gradient around the detection site.

In the previous glucose Sencil, Con A was conjugated with fluorescein isothiocyanate (FITC), which

fluoresces at 520nm when excited at 470nm1. We found that the coupling of small optical fibers to conventional laboratory spectrophotometers produced too much back-scattering of incident light to achieve the desired signal-to-noise values in a single fiber (figure 3A). In the new sensor, a Quantum Dot replaced FITC as the photon donor. The QDot can be excited by a much shorter wavelength than the 470nm required to excite the FITC, making it easier to filter out back-scattered excitation light (figure 2B). Betacyclodextrin is conjugated with tetramethylrhodamine isothiocyanate (TRITC), which fluoresces at 570nm when excited at 520nm (the emission of this particular QDot). The ratio of the two emissions can be used to indicate the concentration of glucose, compensating for fluctuations in absolute fluorescence intensity that tend to arise through variable coupling to the sensor and biodegradation of the matrix over time.

Figure 3: Comparison of spectrums with different photon-donor. (A) FITC-based glucose sensor (B) QDot-based sensor. The truncated peak in both figure to the left are from the back scattering of the shorter wavelength excitation light centered at 470nm (FITC-based sensor) and 420nm (QDot-based sensor).

B A

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2.2 Signal transduction efficiency The Sencil uses a

percutaneous optical fiber to provide secure optical coupling to the sensor, overcoming the unpredictable nature of transcutaneous optical interrogation9-11. In order to permit the fiber to ride freely in the skin when not in use, it is important that it have a diameter and stiffness similar to a natural hair. In order to facilitate insertion through the lumen of a small hypodermic needle, it is important for the sensing matrix to form a compact profile at the end of the fiber similar to a hair follicle. We selected a multimode silica fiber with 100µm core size plus 10µm silica cladding and 10µm polyimide jacket. The small diameter (about twice that of a human hair) offers the potential for use as a continuous monitoring device with efficient photocoupling, easy insertion and minimal irritation to the surrounding tissue1.

Previous photonic chemical sensors attached a hollow tube sealed with a membrane onto the end of the optical fiber12, which is bulky and labor intensive. For the Sencil a roughly droplet-shaped hydrogel made of polyethylene glycol (PEG) is polymerized directly onto the end of the fiber (Fig. 4). The follicle-shaped droplet is formed by shining UV light through the fiber into the precursor solution, which photopolymerizes into the desired shape according to the absorption and dispersion of the light from the end of the fiber. Multiple dip-coats can be used to build up the desired size and shape. The optimal size of the droplet is about 2-3 times the diameter of the optical fiber; increases beyond that provide little additional photonic signal because the excitation light becomes dispersed and any fluorescence it evokes is unlikely to be captured by the small aperture of the fiber. The pore size and hydrophicility of the matrix can be controlled by selecting the molecular weight of the PEG monomer and mixing different compositions and amounts of copolymer to facilitate the diffusion of the analyte. A similar configuration was used to improve the response time of a pH sensor by using a polymer instead of a membrane barrier to control the leaching of sensing materials.13

2.3 Biocompatibility, chemical/mechanical integrity As with any implanted device, strict biocompatibility requirements limit material selection. This can

be divided into effects of the sensor on the tissue and effects of the tissue on the sensor. Because of the intimate contact required between the chemical transducer and body fluids, even reactions that are nominally benign may cause serious degradation of sensor performance. For example, encapsulation by connective tissue may reduce diffusion of analyte to the sensor surface.14 In the worst case, chronic

Figure 4: Manufacture method with adhesion enhancement. Step 1: Surface of the optical fiber is etched with potassium hydroxide and colvalently coated with collagen. Step 2: Dip modified fiber into precursor solution and guide UV light through fiber to induce polymerization. Step 3: Multiple dips and UV polymerization to increase interface area.

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inflammation occurs around the sensor implantation site. Such reactive tissues may have atypical concentrations of analyte and they may accelerate the biodegradation of the matrix polymer15. Despite progress in microminiaturize technologies, conventional implantable devices (mm to cm) remain much larger than cells and blood capillaries (10-100µm). For example, needle-type glucose sensors are small enough to avoid use of anesthetic for self-implantation but their insertion into the subcutaneous tissue creates locally severe damage to cells and capillaries. Smaller implants tend to provoke lesser foreign body reactions, but their small volume may make them more susceptible to damage by leaching or enzymatic degradation of their critical components.16

Optical fibers are small, thin, lightweight, chemically stable, generally biocompatible, immune to electromagnetic interference and inherently electrically isolated, all desirable properties for medical devices. The 110µm silica fibers with 15-25µm polyimide jacket can bear more than 500g tensile force, the point at which our clamps of the force gauge tended to slip. The fibers could be bent repeatedly to a <5mm radius with 36±6g peak force1.

Safety of the PEGs has been studied and reported for over fifty years: protein absorption studies have shown that fibrinogen absorption was significantly reduced after coating polymers either linear or branched PEG;17 photo-crosslinked PEG copolymers induced a much less fibrous encapsulation after subcutaneous implantation in rat.18 The lack of toxicity is reflected in the fact that PEG is one of the few synthetic polymers approved for internal use by the FDA, appearing in food, cosmetics, personal care products and pharmaceuticals.

It is likely that very small quantities of PEG (approximately 10µg) and Con A (<10-12g/sencil; the lethal dose in mice is 50mg/kg by intravenous injection) will be left behind subcutaneously each time a percutaneous fiber is removed. We believe and intend to demonstrate that this PEG will gradually biodegrade without producing adverse or cumulative effects. Histology of chronically implanted sites has verified the biocompatibility of the Sencil.1

The polyimide coat of the optical fiber improve its mechanical strength greatly, but it discourages cell and protein adhesion.19,20 The optical fibers separated easily from the PEG matrix and tended to slip out of

Figure 5: Examination of optical fiber surface with 3 different treatments: column 1 shows a virgin fiber; column 2 was etched by potassium hydroxide; column 3 was etched and covalently coated with Oregon Green conjugated collagen (Invitrogen, Carlsbad, CA), which emits green light when excited at 488nm (bottom row of images). The last group demonstrates the distribution of collagen on the surface. The detail of surface treatment is described in fabrication methods below.

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the skin.1 To address both problems, we decided to bind a layer of collagen covalently to the polyimide coating because collagen is a well-known cell adhesion promoter.21,22 It is also possible to improve the adhesion of the PEG matrix by using NHS-PEG-acrylate as a crosslinking agent (described in fabrication method). In order to investigate the adhesion strength between the sensing matrix and fiber end, a peak-force meter was clamped to the free end of the Sencil. The sensing matrix bulb was pulled until it separated from the fiber. We measured a peak tension force of only 0-2g for untreated 100 µm fibers. It is increased to 40g after collagen treatment and to 70g after cross-linking, which is the mean peak force for removing hair from the human scalp. Chronic stability of the modified Sencils in vivo will be investigated in the near future.

3 Fabrication methods

3.1 TRITC-betacyclodextrin Conjugation All chemicals involving in the conjugation were purchased from Sigma (St. Louis, MO) except TRITC from Invitrogen (Carlsbad, CA). We used the isothiocyanate reaction to conjugate TRITC to betacyclodextrin. Betacyclodextrin is dissolved in methyl sulphoxide containing a few drops of pyridine. TRITC is added, followed by dibutyltin dilaurate, and the mixture is heated for 2 hr at 95C. The TRITC-betacyclodextrin is dried by reduced pressure evaporation to remove solvent.23 The residue is dissolved in phosphate buffer saline (PBS). Dialysis and fluorescence measurements on the dialysate are used to separate unbound fluorophore and quantify the stoichiometry of the bound fluorophore. 3.2 Qdot-Con A conjugation All chemicals in the reaction were purchased from Invitrogen (Carlsbad, CA). 5mg Con A was dissolved in 10ml borate buffer with 20mg succinic ahhydride for 2 hour. Dithiopthrethiol (DTT) was added to reduce the thiol group of Con A with the final concentration at 20mM, and kept in room temperature for 30 minutes. Excess DTT was removed by column chromatography. 4-(Maleimidomethyl) cyclohexanecarboxylic acid N-succinimidyl ester (SMCC) was used to activate Quantum Dot in DMSO with the final concentration at 1mM, and reacted in room temperature for 1 hour. Excess SMCC was removed by column chromatography. The activated Qdot was treated with caprylic acid N-succinimidyl ester. Both activated materials reacted for 1 hour in room temperature. The reaction was quenched by betamercaptoethanol. The Qdot-Con A was separated by column chromatography.

3.3 Adhesion enhancement All the chemicals involved were purchased from Sigma-Aldrich (St. Louis, MO) except 1-ethyl-3-(3-

dimethylaminopropyl)carbodiimide, hydrochloride (EDAC) from Invitrogen (Carlsbad, CA).The optical fiber was treated in 4M potassium hydroxide for 3 hours and washed by pH 5 phosphate buffer. The etched fiber was activated for 15 minutes in a solution with both 50mg/ml EDAC and 50mg/ml N-hydroxysuccimide dissolved in pH 5 buffer. The activated fiber was washed with pH 5 buffer and reacted with 1mg/ml collagen for 2 hours. Then the collagen coated fiber was washed with pH 7.4 PBS and activated with 10mg/ml NHS-PEG-acrylate dissolved in pH 7.4 PBS.

3.4 Sensor fabrication All the fabrication-related chemicals were purchased from Sigma-Aldrich (St. Louis, MO) except for

α-acryloyl, ω-N-hydroxysuccinimidyl ester of poly (ethylene glycol)-propionic acid, MW 3400 (PEG-NHS 3400) from Shearwater Inc. (Huntsville, AL). UV induced polymerization9 is used to attach the biosensing matrix to the internal end of the fiber. In order to immobilize Con A, a trace amount of PEG-NHS 3400 was added into the precursor solution (1/200 mass ratio to PEG-DA) and incubated with Con A for 30 minutes. The NHS ester of NHS-PEG-diacrylate reacted with lysine residue on the surface of Con A to form a covalent bond. Another acrylate end-group of NHS-PEG-acrylate was crosslinked with PEG-DA to form the hydrogel matrix during UV-induced polymerization.

The TRITC-labeled betacyclodextrin we are now using has a molecular weight of 1kD, which is 1800 times reduced in size compared to the molecule that Meadows and Schultz immobilized by trapping in PEG. This small dextran is unlikely to be trapped effectively in the PEG matrix so it needed to be

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covalently bonded to it. We modified the labeled betacyclodextrin with acryloyl group, which provides a covalent binding site for PEG matrix, the same functional group used for the UV polymerization. A solution of acryloyl chloride (0.54g, 6 mmole) in 10ml Ch2Cl2 is added dropwise to a solution of TRITC-betacyclodextrin (3 mmole) and triethylamine (3.2g, 31.7 mmole) in 60ml Ch2Cl2 at –5C over one hour. The reaction mixture is stirred over night at room temperature, and then triethylamine • hydrochloride is filtered off. The filtrate is diluted with 100ml Ch2Cl2 and extracted with 2x50mL NaHCO3 (10%) and 1x50mL brine. The organic phase is dried over MgSO4, filtered and distilled to give crude product.24 The effectiveness of the binding was assayed by measuring the fluorescence of the supernatant after prolonged soaking in phosphate buffer saline.

The precursor solution of the polymer matrix was prepared with 1 mL PEG-DA, 5 mg NHS-PEG-acrylate, and 0.5 ml of 10nM QDot-Con A and 30nM TRITC-betacyclodextrin in PBS for QDot-based sensor (Figure 3B). The previous version (Figure 3A) used 100nM FITC-succinyl-Con A (Invitrogen, Carlsbad, CA) and 300nM TRITC-betacyclodextrin in PBS. After vortexing for 30 minutes, 100 µl of trimethylolpropane triacrylate (TPT) and 10 mg of 2,2-dimethoxy-2-phenylacetophenone (DMPA) were added to the solution, and vortexed for another 30 minutes. This is the precursor solution that is dip-coated and photopolymerized onto the end of the optical fiber (Figure 4). 4 In vitro performance

The in vitro performance of the prototype sensor has been evaluated by measurements in glucose solutions with different concentrations. Figure 3 presents the spectral shifts produced by various glucose solutions. Response time of the device to an increase in glucose was about 5-15 minutes (Figure 6). Fluorescence response time to a step decrease in glucose was about 15-25 minutes. There was no obvious difference in temporal response between the two fluorophore systems. Changing the pore size of the PEG had little effect, suggesting that the limiting factor by the mobility of the Con A with respect to the betacyclodextrin after both are attached to the PEG matrix. The ratio of the two fluorescence wavelengths (Figure 6B) appears to provide a stable but somewhat noisy signal, which can be improved by increasing the concentration of the fluorophores in the matrix and the intensity of the excitation source. The signal has been stable for 5 weeks and the experiment is still underway.

Figure 6: Sensor dynamics. (A) The sensor was place in 0mg/dl glucose solution from 0-12th minutes, and transfer to 100mg/dl from 12th to 49th minutes, 300mg/dl after 49th minute. Green curve indicates the response of Qdot emission, and red curve indicates the response of TRITC emission. (B) The datapoints plot the ratio of emission of QDot divided by emission of TRITC at the corresponding time point.

A

B

[0mg/dl] [100mg/dl]

[300mg/dl]

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5 Discussion and Conclusions

The materials and fabrication methods to produce Sencil implants appear to be robust and inexpensive. The physical and chemical properties of the polymer matrix can be modified to alter its hydrophilicity, crosslinking density, and pore size for specific application in detecting analytes other than glucose. Systematic nano-engineering of binding affinities, fluorophore attachment, reactant entrapment and matrix structure provides the key to a wide range of clinical applications. We now have a variety of chemical tools with which to optimize systematically the nanostructure of the sensing matrix. Methods to improve adhesion between the matrix and the optical fiber are promising but remain to be demonstrated in vivo. The photonic integrity and chemical stability of the sensing matrix also need to be determined by serial photonic measurements over the design requirement of 1-3 months in vivo.

Ideally, the sensor will compute the glucose concentration from the ratio of the two emissions. For the sensor prototype, the fluorescence spectrum (figure 3) indicates that the two emission spectra are overlapped, so that the peak amplitudes of each are influenced by the skirts of the other. More accurate measurement can be achieved by estimating the area-under-the-curve attributable to each fluorophore by using independent component analysis (ICA).25,26 ICA maximizes the degree of statistical independence among outputs using contrast functions approximated by the Edgeworth expansion of the Kullback-Leibler divergence. In contrast with decorrelation techniques such as Principal Component Analysis (PCA), it ensures that output pairs are uncorrelated. ICA imposes the much stronger criterion that the multivariate probability density function (p.d.f.) of output variables factorizes. Finding such a factorization requires that the mutual information between all variable pairs go to zero. Mutual information depends on all higher-order statistics of the output variables while decorrelation only takes account of 2nd-order statistics. ICA combined with enhanced signal strength should provide the accuracy and precision required for a clinical glucose sensor.

The next stages of research will involve chronically implanting Sencils in animals to demonstrate mechanical stability and allow comparison of measured interstitial glucose concentrations with simultaneous blood glucose measurements using laboratory equipment. If the technology appears promising for glucose, we will explore the performance limitations of photonic technologies suitable for miniaturization into a handheld, battery-powered detector. Reference

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22. Jokinen, J., Dadu, E., Nykvist, P., Kapyla, J., White, D. J., Ivaska, J., Vehvilainen, P., Reunanen, H., Larjava, H., Hakkinen, L., Heino, J. “Integrin-mediated cell adhesion to type I collagen fibrils.” J. Biol. Chem. 279: 31956-63, 2004.

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