Patent Application: US-78566007-A

Abstract:
the present invention provides a novel biosensor and measuring method . the novel biosensor of the present invention comprises an electrode having a nanopore structured and catalytically active cyclodextrin attached thereto . the biosensor of the present invention has demonstrated robust analytical performance for direct glucose measurement without mediators or without using native enzyme , which is especially beneficial in the hypoglycemia range .

Description:
reagent grade poly ( 4 - vinylpyridine ) ( pvp ), polyethylene glycol diglycidyl ether ( peg ), triacetyl - β - cd ( t - β - cd ), β - cd / epichlorohydrin , β - d - glucose were purchased from aldrich - sigma . the pvp was recrystallized in methanol . the biomimetic glucose enzyme , which is a biomimetic histidine residue ( his - 516 ) receptor of glucose oxidase and mimics the active center of native glucose enzyme , named mm - β - dmcd was synthesized generally according to the published procedures ( e . t . chen and h . l . pardue , analytical applications of catalytic properties of modified - cyclodextrins . anal . chem . 65 , 2563 - 2567 , 1993 , which is hereby incorporated by reference in its entirety as if set forth herein ). u . s . pat . no . 6 , 582 , 583 issued on jun . 24 , 2003 is also hereby incorporated by reference in its entirety as if set forth herein . briefly , β - dmcd may be reacted first with sodium hydride in dry tetrahydrofuran under a nitrogen atmosphere at 35 - 38 ° c . for 10 hours . the solution is then cooled to 0 ° c . and mixed with a solution of 2 -( 4 - imidazolyl )- ethyl bromide in tetrahydrofuran and heated to 25 ° c . for 10 hours to produce the mm - β - dmcd . the structure of the mm - β - dmcd is shown in fig1 . a gold electrode ( 1 . 6 mm diameter ) polished successively with 0 . 1 and 0 . 05 μm alumina slurry ( bas ), then washed with double distillation water , then sonicated with methanol , then with water . after that , the electrode was polished with diamond solution ( bas ), and washed with double distillation water and sonicated in methanol , then with double distillation water . dry n 2 was used to dry the electrode , and then the gold electrode was put in a 35 ° c . incubator for further drying for 1 hour before use . the gold electrode with a sam film was used as the working electrode . the platinum wire electrode was the auxiliary electrode and the ag / agcl electrode was the reference electrode . a class 100 level of a clean room was used for all sam developments . a mixture of pvp / peg / mm - β - dmcd ( see e . t . chen . amperometric biomimetic enzyme sensors based on modified cyclodextrin as electrocatalysts , and u . s . pat . no . 6 , 582 , 583 issued on jun . 24 , 2003 , both of which are hereby incorporated by reference in entirety as if set forth herein ) solution ( e . g . 4 μl ) was dropped using a syringe by 2 × 4 μl onto the gold electrode surface at a room temperature and the fabricated sam electrode was immediately sealed in a n 2 filled container and incubated for 48 hours at 35 . 0 ° c ., then the electrode was washed with double distilled water to remove unbounded chemicals , then was incubated for 2 hours before use . the same protocols were used for fabrication of the peg / mm - β - dmcd sam film without pvp ; and a t - β - cd / peg / pvp / β - c copolymer sam sensor was also fabricated under the same procedures . the differences in the composition and concentration between the u . s . pat . no . 6 , 582 , 583 and an embodiment of the present invention is shown below : it should be noted that different factors have impacts on the formation of different nanostructured sam film on a gold surface . a comparison of these factors in an embodiment of the present invention and u . s . pat . no . 6 , 582 , 583 is shown below : a single crystal gold 1 × 1 × 1 film that causes phase structure transition was reported in y . kondo et al . ( see reference 28 ). the different thickness of the gold film has an impact on the formation of the sam film on the gold surface . in addition , according to u . s . pat . no . 6 , 582 , 583 , the gold planer electrode was immersed in the solution for 24 hours at a room temperature . however , in an embodiment of the present invention , only one drop of the solution was applied onto the gold chip surface . after the application , the solution was immediately taken into incubation . the step of immersion in a sealed temperature for 24 hours at a room temperature was skipped . a clean bare gold chip with 50 nm thickness and 3 mm diameter was purchased ( genefluidics , ca ) for fabrication of the cd - sam . pretreatment of the chip before the fabrication is not necessary based on the afm image of the bare gold surface . the same procedures and chemical mixtures as above were used to fabricate the gold cd - sam chip in the clean room for the afm measurements . the morphology of the three cd - sams against a bare gold electrode was characterized by using an instrument ( digital instruments dimension 3100 atomic force microscope , veeco instruments , santa barbara , calif .). the nanopore sizes were measured using tappingmode ™ afm with a silicon cantilever and tip with a 300 khz resonance frequency and a 5 - 10 nm tip radius ( model tesp by veecoprobes ). the software used was nanoscope versions 5 . 30rl . the first reported nanopore structured biomimetic cd - sam was shown in fig1 a ( two dimensional view , roughness measurement ), fig1 b and 1c ( 3d view ), fig1 d ( illustrative drawing ) and fig1 e ( pore size measurement ). the images clearly revealed the smoothness of the sam and the fact that the nanopores were evenly distributed and vertically oriented on the gold surface with the pore size from 10 nm to 20 nm , and the roughness of the sam was 0 . 82 nm rms . fig1 d and 1e show the example of the pore size of 19 . 5 nm . fig2 a and 2b are the 3d afm images for a sensor with the same chemical composition and the receptor , but without nanopore structure . the nanopores were not observed . however , a “ forest ” of nano pillars ( 10 - 60 nm diameter ) was observed covering the gold surface with a relative roughness of 16 . 65 nm in the z direction of the membrane , which was much rougher than the former sensor . fig3 a and 3b are the afm images for another type of sensors that were fabricated by the inventor , which had the same configuration as the sensor in fig1 a , except that triacetyl - β - cd ( t - β - cd ) instead of the receptored cd was used . the relative film roughness of the sam membrane was 24 . 6 nm , which was too rough and the signature nanopore structure was not observed . a voltammetric analyzer ( model cv50w , bioanalytical system ( bas ), in ) was used for the measurements of currents . a faraday low current cage ( model c2 , bas ) was used for protection of the electrode cell . for the ph effect study and for the glucose measurements , the scan rate was kept constant at 50 mv / s . all electrochemical measurements were done in an unstirred electrochemical cell at 20 ° c . all sample solutions were bubbled thoroughly with high purity n 2 for 10 minutes and maintained in a n 2 blanket . the 0 . 1 m , ph 7 . 0 ± 0 . 1 buffer (( 0 . 1 m kcl ) solution was filtrated and degassed . the electrodes were equilibrated in a 10 ml , ph 7 . 0 ± 0 . 1 , 0 . 1 m buffer ( 0 . 1 m kcl ) for 30 - 45 minutes by applying a potential at − 400 mv until a steady - state current was observed before a sample can be measured . the internal standard addition method was used to study the accuracy of glucose measurements using bovine serum albumin ( bsa ). the current for a 50 mg / dl glucose standard was measured in the 0 . 1 m phosphate buffer , ph 7 ( 0 . 1 m kcl ) bovine serum albumin . then 100 μl of 5 g / dl of glucose solution was added into the sera , and the current was measured . four measurements were obtained after 4 consecutively additions of the same amount of glucose solution . the electrochemical behavior of the sensors was characterized by using cylic voltammetry ( cv ) method . the factors affecting the currents were studied . the cyclic voltammograms of different electrodes with and without nanopore structured sam membranes are compared in fig4 a , 4 b and 4 c . in fig4 a , a well - defined irreversible reduction peak was observed for the nanopore sensor curves a and b , indicating that the nanopore structured cd - sam was favorable for the det between the active center of the imidazolyl in the cavity of mm - β - dmcd and the electrode . the decrease of the current shown in curve b indicates that the glucose molecules entered the cd cavity and mingled with the active receptor , hence suppressing the det between the receptor and the electrode . fig4 b shows the electrochemical behavior for the t - β - cd &# 39 ; s sam electrode . the curves a and b have large envelop background currents . no det peaks were observed for the bare gold electrode and for the t - β - cd electrode . fig4 c shows that there is no det peak for mm - β - dmcd without nanopore structure , even it has the mimic his receptor , in the presence or absence of glucose . in fig4 c , the curves a and b overlap and the heavy envelop - like background currents exist , which was consistent with the morphology of the afm image . fig4 a shows the electrocatalytic current and fig4 c does not have the current , even both sensors had the same biomimetic receptor , the differences being that the biosensor in fig4 a has the nanopore structure and the biosensor in fig4 c does not have the nanopore structure . this indicates that a lack of nanopore structure could hamper the det even in the presence of an active receptor . the scan rate effects on the electrochemical behavior of the nanopore cd sensor were studied and the voltammogram profiles were shown in fig8 . the reduction peak currents increased as the scan rate increased in the studied range from 20 mv / s up to 450 mv / s . the linearity study of the scan rate effect on the e p , c values is presented in fig9 . the nanopore structured cd sensor distinguished itself from other reported sensors that had reversible redox peaks ( see references 3 , 9 , 25 ) and associated with the det effect , which was the irreversible direct electron transfer . possible explanations were that the effects of the nanopore structures were significant on det . it played a significant role in promoting the det . according to the commonly used e . laviron &# 39 ; s method , the det rate constant for one nanopore structured cd - sam sensor was calculated as 131 ± 2 . 3 / s based on three replicate measurements in neutral buffer , which had a 3 . 4 - fold increased det compared with 38 . 9 ± 5 . 3 / s for the rate constant for a gold nanoparticle - based glucose sensor using native glucose enzymes ( see reference 9 ). the results also had a 3 . 11 - fold faster rate than a god glucose sensor with single - walled carbon nanotubes ( see reference 12 ). some of the advantages of the nanopored cd sensors of the present invention over the prior art native glucose enzyme sensors with gold nanoparticles or carbon nanotubes are : ( 1 ) the activation of the biosensor without the need of the presence of oxygen to detect glucose simplifies the procedures for commercialization ; ( 2 ) the fabrication of truly reagentless , mediatorless nanopore cd sensors without the use of glucose enzyme avoided biofouling and denaturing from using native enzymes , which is an attractive characteristic for implantable devices or for usage in harmful environments . the change of ph effects on the electrochemical behavior of the nanopored cd sensors was evaluated in 0 . 1 m phosphate buffer with varied ph from 5 . 0 to 9 . 0 without the presence of glucose at 20 ° c . as shown in fig5 . the highest peak intensity was observed at ph 7 . 0 . the cathodic peak diminished at ph 9 . 0 indicated more negative ions from the solvent solution suppress the det electron flow . the peak shifting slightly to a positive potential due to a decrease of ph was also observed . therefore , the sensor is useful over a ph range of from about 5 . 0 to about 8 . 0 . under optimal experimental conditions , curve c in fig5 shows the optimal results , where det occurred at a reduced potential around − 390 mv . the cyclic voltammogram profiles are shown in fig6 upon the addition of various standard glucose concentrations successively in the 10 ml ph 7 . 2 buffer solution . as shown in fig4 a , for curve b , the current decreases in the presence of glucose . the fact that electrocatalytic current increased proportionally with higher glucose concentration indicates that the channeling effect due to the nanopore structure had overcome the effect of glucose - receptor reaction resulting in the temporary suppression of the direct electron transfer . recent published literature has revealed the fact that a decrease in current was observed as analyte concentration increased in gold nanoparticle sensors when native enzymes were used ( see references 3 , 9 ). this further provided evidence proving that when β - cd is lodged in the lumen of the α - hemolysin ( hl ) pore , it reduces the unitary conductance by about 70 % ( see reference 16 ), and the current reduces significantly when a voltage is applied onto the biological system in comparison with a system without an β - cd entering the α - hl pore . the experiments of the present invention not only confirm the nanopore sensor &# 39 ; s electrochemical function , but also reveal a distinct phenomenon : at the beginning , a decrease of current is due to the association of the glucose molecules with the receptor site , and after that , an proportional increase of current is due to the nanopore channeling effect when the glucose concentration continues to increase . detailed illustration of the pathway of the nanopore sensor is presented in fig1 . for within - run precision , the relative mean standard deviation ( rsd ) was 1 . 5 % from the triplicate runs obtained at each of 11 glucose concentration levels from 5 to 100 mg / dl . at the clinical decision level of 50 mg / dl , the rsd values were 1 . 1 % and 1 . 4 % ( n = 5 ) obtained at different days using the same nanopored cd sensor # 1 . at 20 mg / dl , which is a useful clinical decision level for diagnosing type i diabetic in newborns ( see reference 23 ), the rsd value was 1 . 5 %. for the inter - assay precision , the rsd values obtained from three cd sensors # 1 with the same nanopored fabrication were 1 . 1 %, 0 . 7 % and 2 % at 50 . 0 mg / dl glucose concentration with five replicates . the precision measurements of glucose at hypoglycemia range from the nanopored cd sensors have laid a foundation for accurate performance for future glucose monitoring devices . this improvement of the analytical performance has overcome the disadvantage of imprecise measurements common to self - monitoring blood glucose ( smbg ) devices of the prior art at the low glucose range ( see reference 23 ). three same types of nanopore structured cd sensor were fabricated on three 1 . 6 mm diameter gold electrodes and were used for the reproducibility study . the det rate constants can be reproducibly obtained . the k s value was 136 . 7 / s ± 19 / s . the peak intensity deviation among the three sensors was 7 . 7 %. the internal standard addition method was used to study the accuracy of glucose measurements using bovine serum albumin ( see reference 26 ). four measurements were obtained after 4 consecutively additions of the 100 μl of 5 g / dl of glucose solution into the bsa . the results were compared against an internal standard . the mean accuracy was 98 %± 1 % at 50 mg / dl concentration . in prior art , native glucose enzyme sensors can suffer biofouling in which the glucose enzyme is easily dissociated from the electrode surface ( see reference 9 ), and , therefore , it needs constant enzyme activity renewal in a solution . this problem does not occur with the nanopored cd sensor of the present invention . the cd sensor of the present invention never needs such a renewal process and still maintains a good performance . for example , the intensity of the same cd sensor only decrease by 16 % after 116 measurements lasted for 42 days . plus , the sensor does not need to be kept at 4 ° c . for storage as required by native enzyme sensors ( see reference 3 ). therefore , the nanopore cd sensors of the present invention have offered advantageous features that are simple and robust for direct glucose measurements without using glucose enzymes or mediators . as shown in fig6 , the well - defined electrocatalytic response curves for glucose are presented by utilizing the arrayed - nanopore sam with an artificial electrocatalytical functioning receptor . a plot of current vs . glucose concentration illustrates the linearity of the nanopored cd sensor &# 39 ; s analytical performance presented in fig7 . the least - squares statistical results obtained from current vs . glucose concentrations produced an equation y ( na )=− 0 . 9 ( na )+ 1 . 97 ×( na / mgdl − 1 ) with a linear range up to 205 mg / dl with the correlation coefficient of r = 0 . 998 , s y / x = 10 . 7 na . the sensitivity of the sensor is 3 . 55 na / μmol / l in 2 . 01 mm 2 electrode surface , which is 118 - fold sensitive than that of the prior arts ( chen , 2003 , see reference 17 ), and 33 . 040 - fold enhanced the sensitivity compared with liu &# 39 ; s glucose electrochemical cyclodextrin polymer sensor ( liu et al . 1998 , see reference 27 ). the calculated limit of detection ( lod ) for glucose using the current invented arrayed - nanopored sensor is 3 . 1 nm / mm 2 , which are 1 . 9 × 10 3 molecules of glucose / nm 2 . this glucose biosensor of the present invention demonstrates the full usages of monitoring glucose at critical clinical decision concentration ranges ( fig7 ) from hypoglycemia to hyperglycemia ranges . the least - squares statistic result in the hypoglycemia range from 5 to 50 mg / dl produced an equation of y =− 0 . 008 μa + 0 . 007 ×( μa / mg / dl ) with correlation coefficient of 0 . 999 ( n = 30 with three replicates at each of 10 concentration levels ), and has the s y / x value of 0 . 006 μa , corresponding to a relative standard deviation of 1 . 6 % at the 50 mg / dl clinical decision level for type i diabetic hypoglycemia . in addition to gold , glassy carbon can be used for construction of the biosensor of the present invention . the det effect was observed and the irreversible peaks were also obtained . the foregoing is considered as illustrative only of the principles of the invention . further , since numerous modifications and changes will readily occur to those skilled in the art , it is not desired to limit the invention to the exact construction and operation shown and described , and , accordingly , all suitable modifications and equivalents may be resorted to , falling within the scope of the invention . 1 . l goorton , a . lindgren , t . larsson , f . d . munteanu , t . ruzgas and l gazaryan , direct eletrontransfer between heme - containing enzymes and electrodes as basis for third generation biosensors , anal . chim , acta 400 , 91 , ( 1999 ). 2 . s . j . updike , g . p . hicks , nature 214 , 986 , ( 1967 ). 3 . z . dai , f . yan , j . chen and h . ju , reagentless amperometric immunosensors based on direct electrochemistry of horseradish peroxidase for determination of carcinoma antigen - 125 , anal . chem . 75 , 5429 , ( 2003 ). 4 . y . tian , l . mao , t . okajima , t . oshaka , anal . chem 74 , 2428 - 2434 , 2002 . 5 . t . ruzgas , e . csöregi , j . emneus , l gorton , and g . marko - varga . anal . chim . acta , 330 , 123 , ( 1996 ). 6 . s . y . lu , c . e . li , d . d . zhang , y , zhang , z . h . mo , q . cai and a . r . zhu . electron transfer on an electrode of glucose oxidase immobilized in polyaniline , j . electroanal . chem . 364 , 31 , ( 1994 ). 7 . p . de taxis du poet , s . miyamoto , t . murakami , j . kimura and i . karube . direct electron transfer with glucose oxidase immobilized in an electropolumerized poly ( n - methylpyttole ) film on a gold microelectrode , anal . chim . acta , 2365 , 255 , ( 1990 ). 8 . j . wang , l . liu , l . chen and f . lu , highly selective membrane - free , mediator - free glucose biosensor , anal chem 66 , 3600 , ( 1994 ). 9 . s . liu and h . ju , reagentless glucose biosensor based on direct electron transfer of glucose oxidase immobilized on colloidal gold modified carbon paste electrode , biosensors and bioelectronics , 19 ( 3 ), 177 , ( 2003 ). 10 . j . zhao , r . w . henkens , j . stonehuemer , j . p . o &# 39 ; 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