CA2080022C - Implantable glucose sensor - Google Patents

Abstract

Implantable enzymatic sensors (25, 43, 44) for biochemicals such as glucose are provided having an ideal size and geometry for optional long term implantation and linear responses over the concentration ranges of interest. The sensors (25, 43, 44) include an elongated body (10, 26, 46) supporting an indicating electrode section having an appropriate enzyme immobilized thereon to present an enzymatic indicating surface (21, 33, 54). A permeable synthetic polymer membrane (24, 42, 56) is applied over the sensor body (10, 26, 46) to protect the enzyme and regulate diffusion of analyte therethrough, to ensure linearity of sensor response. The sensors (25, 43) are of flexible design and can be implanted using a catheter. Alternately, the sensor (44) includes an internal indicating electrode body (46) housed within an apertured, hollow needle (48). A holder (66) affixed to the needle (48) allows for easy manipulation and implantation of the sensor (44).

Description

wO 91/15993 PCI/US91/02641 IMPLANTABLE GLUCOSE SENSOR
This is a Continuation-ln-Part of Application S/N 07'511.059. filed April 19. 1990 Back~round o~ the Invention a 1. Field of the Invenlion The present invention is broadly concerned with a subcutane-ously implantable enzymatic sensor charac~erized by small size. oplimum geometry and linearity oE sensor response over the concentration range of interest. More particularlv. it is preferablv concerned with an imFl~nt~hle glucose sensor oE this type designed ~o p~ovide, in conju~,lion with a suitable signal processing unil. a currenl which is propor~ional ~o subcutaneous glucose concentration. In preferred forrns. glucose sensors of the inven~ion are based on the en_yme-cataly_ed oxidation of glucose ~o gluconic acid and hydrogen peroxide, ~he latter being moni~ored a~ e,o,l.c~rically by the sensors.

2 D~~ ion of ~he Prior Art There have been a great many a~tempts in the past to develop t~iable rl_J~?~'~ sensors for ~.,Lin,lous in vivo ~ ,asu.~,...en~s of bio, h~...;-cals. For example. conside.àblc effort has been made to devise reliable implantable sensors ~or ,..onito,i..g glucose concen~.dtions in blood. Such determin~fior~c are useful in a variety of applications, e.g., in the ~.~,.. ;.. ~"t of irc One difficulty in pl~ ing a reliable ~ nt;-l''S glucose sensor is that glucose levels in the bloo bllcal.. of a patient vary on a time basis and are normally ~ pe ~d ~ upon the physical activity oE the individual. his food, beverage and sugar intake. his m~t~bolir ra~e, and o~her ind;~ i7Pd fac~ors.
Furthermore, the geo....... ,l.~ of the sensor must be such as ~o adapt to implanta-fion in a living patient.
Glucose sensors have been p.ul osed in ~he past which rely upon the well-r~ctablichr~d enzyme-catalyzed ~Yi.1~tic n oE glucose wherein glucose and oxygen function a. substrates ~or the enzyme glucose oxidase in the production of gluconic acid and hydrogen peroxide, ~he latter being l.. o-.ito.cd ampero-metrically. See, for example, U.S. Patents No. 3,539,455 to Clark and 4.671.288 ~ to Gough.
Although the idea of an imrl~n--\-'s ellL~IIId~iC glucose sensor is not ~ new, cor~iderab'~ difficulty has been el.co~,nte,~d in p~udu~hlg reliable.
cost-efficient devices o~ this character. For ~Y~ , many p.opos~ sensor . ~
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3 Pcl/US91/o2641 2 ~ 8 ~ 0 2 2 geometries are simply not re~lictic~lly implantable. at least for the periods of time required for adequate clinical glucose monitoring. Thus, the devices proposed in the '288 Gough Patent, because of a requirement of multiple elec-trodes carried within a tubular needle, inevitably are of such diameter as to beS uncomfortable to the user and not practical for extended iml l~nt~tion.
Furthermore, many prior sensors do not exhibit a stable and linear response, particularly over PYt~nAed times of impl~nt~tinn, and do not give accurate and reliable results. Finally, fa~.icaLioll of prior glucose sensors has presented formidable difficulties, to the extent that only about one in five sensors pro-duced by ~,on~cnlional te.,l.n;~ues are deemed usable. This obviously repre-sents a considerable inefficiency, to the point that no truly successful implant-able glucose sensor has helel~role been p.uduced on a large scale.
Summar~ of the Invention The present invention o~e.~lllLs the plùbl_...s outlined above, and 1~ provides a greatly i--.~J~u.ed ~,.l~lllal;c sensor specifically designed for long-term impl~nt~tinn in a patient. The sensor is adapted Eor pncitinning in an em,hc l".,cnt characterized by the presence of biological mr~ ec which are substrates for or products y~os~uc~d by enzymes, in order to d~rmin~ the presence of such biological molecules. While the p.h~ci~lcs of the invention may be used in the fabrication of glucose sensors, the invention is not so limited. Indeed. the sensors in a~u,dancc with the invention may be produced using a wide variety of h..-..obil;~d enzymes. for the detection of an equally large number of analytes. Exemplary enzymes and their cullc~yonding sub-strates are given in U.S. Patent No. 4,721,677 to Clark, and this patent is incorporated by reference herein.
In any event, the enzymatic sensors in accordance with the invention preferably are in the form of an elongated body supporting at least an indicat-ing electrode, with the inrlir~ing electrode presenting a seclion adapted for exposure to the biological environment. The inr~ir~inv electrode section has an enzyme operably immobilized thereon to present an enzymatic inrlir~ing surface. A number of variants are possible for the reference electrode. For PY~m?la~ use may be made of an exteïlialiy applied eleciiocd~diog~ai" ~ic;"
electrode (an 8 mm disk covered with silver chloride and available as Model E-24~ frûm the Phymep Company, 21 Rue Campoformio, Paris, France)~ or a reference electrode which is implanted with the in(licating electrode.

WO 91/15993 PCl/US91/02641 3, 2080022 In one specific embodiment employing an implanted reference elec-trode, the indicating surface oE the irl-iir~ting electrode and the reference elec-trode are laterally spaced apart along the length of the body and each substan-tially ~h~,ulll~libes the latter and is sllhs~:lnti~lly exposed to the biological environment when the sensor is placed therein. Use oE such circumferentially f Ytentling enz~matic in~irating surfaces and reference electrodes sections is believed to be an important aspect of this embodiment. Alternately, the reference electrode section may c~ ..ise a CUIJdU~ Ve salt bridge ~ .Ulll~
ing the body and Iying in a plane transverse to the lon~itl.-iin~l axis of the body; in this case, a lcr~,~cnce electrode is placed in electrical contact with the salt bridge, through use oE a buffered electrolvte. In another embodiment, the reference electrode is simply placed adjacent the inrlirating electrode as a part of the overall sensor.
In preferred practice, the sensor body advantageously comprises an f 1~ lly COf~lu Li-,e noble metal ~e ~ pla~num or ~ iridium) electrode covered with electrically insulative material. with a portion of this material removed from the ele~ ,dc to define an enzyme--ccelvil~g zone. Thus, a short length of Teflon (polytetrafluoroethylene) coated platinum-iridium wire may be provided, with a short section of the inclll~tion removed intermediate the ends of the wire, so that respective segmt ntC of the incul~ting material are on opposite sides of and define a recessed enzyme-receiving .,i~culllfe~c.~lial zone.
Alternately, the endmost portion of the Teflon may be removed, leaving a protruding exposed stretch of wire which defines the enzyme-receiving zone.
An enzyme is operably immobilized on the exposed section of the platinum-iridium wire~ by known means such as adsorption of the enzyme on a cellulose acetate or Nafion layer (1-3 microns thirL-nf~cc), followed by cross linking with giutaraldehyde.
Another important aspect of the present in-ention resides in the preferred use of a synthetic polymer membrane disposed over the enzymatic in(iic~ting surEace to serve as a permeable protective layer. In particular. a layer of polyurethane is advantageously applied as a thin coating over at least thc inr~ ng su.facc (and ;;._.~ra'lj- thc clliirc inr1icat;np electroLlc) hi oruc~
to protect the el"yll~atic reaction surface from the biological environment.
Moreover, this layer provides a diffusional barrier Eor glucose which slows down3~ the flow of glucose and creates a linear sensor response over the concentration Wo 91/15993 PcltUS9l/o264 2 0 8 0 ~ 2 2 ranges of interest. In particular, in order to achieve the desired linear re-sponse, use is made of an active enzyme layer and a relatively thin protective membrane. It is important that the membrane regulate the passage of mole-cules therethrough to an extent that the enzymatic reaction between the in-iirating sur&ce and these molecules is determined by the rate of diffusion through the Illu~llblanC, and not the en_ymatic reaction kinetics. In practice using the methods of sensor construction herein described, an optimal balance between the ~ el;~& goals of linear response and se.~ ivily and response times may be achieved.
The use of an ad~ ;n~ negatively charged inner l"~,.. ,b,~.ne layer immlorliately adjacent the Pt-lr wire also retards the diffusion of negatively charged species (e.g. ascorbate and urate) in the biological e..vi.olll..-,..t which are interfering species. Of course, this inner ~ .,.b~dne does not cignifir~ntlyexclude hydrogen peroYide, an electrically neutral species.
Although the l1.~ ~ of the o~ l pol~e~l~an~ ' .......... ne has not been spe~ifir~lly ascertained, it is ~ctim~ted that the membrane has a of from about 5 to 10 microns in the p~f~ ;d glucose sensors hereof.
The sensors described above are, by virtue of their construction, relatively fleYible and therefore comfortable in use. However, this same cll&~ liaLic Ele,Yibility makes it nccessa,y to employ a catheter to implant thesensors. In an alternative embodi".~nt, sensors may be provided which can be readily implanted without the need of a catheter, even by the patient himself.
In such embodiments, use is made of an Plong~t~l tubular, metallic housing, typically a conventional hypodermic needle; the sensor apparatus is inserted within the needle, and includes an inrlirating electrode having a section thereof provided with immnbili7.od en_yme. In order to expose the en_yme to the biological environment, the needle sidewall is apertured in registry with the en_yme. A holder is also provided adjacent the rearward end of the needle body in order to facilitate manipulation and insertion of the sensor. This holder advantageously is in the form of a transversely PYtPn~ling flag-like plastic bv '; ..ccu,~d to the needle housing.
The invention also collll),ehel,da a novel method of applying the polyurethane membrane described previously. That is to say, a real difficulty inthe production of enzymatic sensors stems from the difficulty of applying .

WO 91/15993 PCl/US91/02641 2~8~022 various materials uniformly to a very small, implantable device. rnis difficultyhas been overcome in the context of the present invention, by applying to the sensor surface a well-defined volume of a polymer dissolved in an organic solvent such that lhe film is unitormly distributed across the surface. In S practice, this method is carried ou~ by providing a wire loop, and holding the coating liquid in the loop by sur&ce tension to form the desired polymer solu-tion droplet, followed by passing the electrodes through the loop to achieve uniform coating along the length of the sensor body.
The enzymatic sensors of the invention have an ideal geometry for imp!~ a~jon Generally spe~kingl the flexible units not housed within a needle are equivalenl in size and shape to a 26-gauge needle (i.e., about 0.45 mm.
oulside diameter). Moreover, ~heir geometry perrnits the reproducible deposi-tion of films and materials and allo vs careful control of the amount and orien~tinn of the enzyme onto the in~ a~ing electrode. hnally, the preferred ser~sors are e~f~Li~ "capped~ with ;~ n~~ (Te~n) which ~revents the sensors from penetrating further into the tissue than is required. Thus. the h~ ion of the sensor causes minimal trauma to the tissue and to the sensor itsel~ The sensor can flex laterally, and this again ...i~ tissue damage caused by movement of the patient.
In the case of implantable glucose sensors, response times of less than two minutes and linearities over glucose concc,nt~ations of 0-25 mM can be achieved. At the same time, through use of the fabrication ~C~ l~n;~lUf ~ of theinvention, the rejection rate upon initial m~nuf~t~lre is drastically reduced.
In the case of sensors received within a needle housing, such can be readily manipulated by the patient for implantation pUIl~OSCS. These sensors typically have a slightly larger diameter than the flexible sensors described previously, but are not so large as to cause ~ ;cco...folL. This relative-ly small size is assured because of the sensor construction, making use of a small Teflon-coated Pt-Ir wire and immobilized enzyme.
Brief De~iv ~c of the Druwin~s Figure I is an enlarged, sectional view illustrating a glucose sensor in accord~ncc; wiil~ e inv~nli~n;
Fig. 2 is an enlarged, sectif n~l view o~ another glucose sensor in accordance with the invention;
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WO 91/15993 PCI~/US91/02641 2 0 8 0 ~ 2 2 Fig. 3 is a graph showing the linear sensor response of the Fig. 1 glucose sensor over a glucose Col1CC~llrdtiOn range of 0-25 mM;
Fig. 4 is a graph illustrating the storage stability of the Fig. 1 glucose sensor;
Fig. 5 is a sectional view depicting another sensor embodiment wherein the inllicating electrode is housed within an impl~nt~hle needle:
Fig. 6 is a perspective view of the sensor illustrated in Fig. 5; and Fig. 7 is a p.,.S,~,~.,livc view of an emho~limpnt similar to that of Fig. 6.
but depicting the use of an implantable reference electrode.
D~ t- of the ~lef~ A ~,d F
The following ~.dlllples illustrate the construction of glucose sensors depicted in Figs. 1 and 2~ and are d~crrihed with particular ~c;Lelt,nce to the,se drawings. It will be understood. however, that the .oY~mplP5 are illustrative only, and nothing therein,should be taken as the limit~ti~n UpOli the overall scope of the i~.. ~
Example 1 - Fi~. 1 One end of a 10 cm section 10 of Medwire Co"~o~ation Teflon-Coated platinum-iridium wire is provided. The section 10 includes a central pl~tinllm_ iridium wire 12 (0.18 mm o.d.) and a coating of insulative Teflon 14 (0.035 mm shirlrnPc~) the.Garou"d. The central wire 12 forms the int~ir~ting electrode from the sensor. A cavity 16 (1-3 mm in length) is formed in the wire 10 as shown in Fig. 1. This is achieved by first putting a circular cut on the Teflon coating with a paper cutter and then pulling the Teflon out to create a cavity of about 1 millimf~ter in length, exposing a corr~pQn~ling section of the wire 12. The excess Teflon extending beyond the left end of the wire 12 is then trimmed off with the cutter.
The ~ ce electrode 18 is formed on the Teflon surface, about I.5 millim.oters to the right of thé exposed platinum iridium surface as viewed in Fig. 1. A thin silver wire (0.1 mm o.d., 15 cm length) is tightly wrapped aroundthe teflon surface covering to form a coil 20 of about 5 millimt~t.ors in length.
A wire wldpping tool may be utilized for this purpose. The trailing portion of the wirc ;u ih ;; liglli Or coil 20 i~ /VCICl wiin a section 22 of heal shriJI~cabi~
Teflon tubing (5 cm long, 1.5 mm o.d., Zeuss Industrial Products Inc.), leaving small lengths of the silver wire and platinum iridium wires uncovered to serve as electrical leads. A heat gun operating at 600~C is employed for shrinking , .

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WO 9l/15993 PCltUS91/02641 -7- 2~0022 !j . '.j.
the Teflon tubing. A layer of silver chloride is formed on the coil 20 by passing current (0.4 mAlcm2) for 60 minutes through the wrapped silver wire while it is dipped in a stirred ().1 N HCL solution. The exposed por~ions of reterence electrode 18 are then rinsed with de-ionized water for 6 hours. The reference electrodes plepaled in this manner show a potential of -64 + 3 mV
(n=10) vs. Ag/AgCl(3M NaCl) in 0.15 M NaCl at 370C.
In order to immobilize glucose oxidase (GOx, E.C.1.1.3.4) on the exposed portion of wire 12, an inner, negatively charged l~ lblane is first applied to the exposed wire section. Thereaher. a circumferentially ~Y~.on~ling erlzymatic inrli~,q.~ing layer 21 is formed within cavity 16. Two different ap-plUd~,ll(.,S have been employed to achieve these ends.
Attqrhm~nt of GOx to bovine serum albumin coupled cellulose acetate The exposed platinum iridium surface within cavity 16 is degreased by washing with acetone. It is then rinsed with de-ionized water and dried in cold air stream before polymer deposition.
The left hand part of the sensor (portion to the left of the ~ ,nce electrode coil 20) is dipped into 5% cellulose acetate (39.8% acetyl content) inSO~o acetone and 50% ethanol for 10 seconds and is withdrawn slowly. It is 2~ then exposed to the vapor above the cellulose acetate solution for 5 seco.. ds and is dipped again into the cellulose acetate solution for 10 seconds. The sensor is then removed and dried in air at room temperature (230C) for one minute and placed in d.,;olliGed water for 6 hours to permit disp!qrPm~n~ by water of entrapped solvent in the llle.llbldne pores. The cellulose acetate Ill~lllI"dnc prepared in this fashion shows good long-term stability and also discriminates well against ascorbate and urate. Bovine serum albumin (BSA) is then covalently coupled to cellulose acetate and a subsequent reaction of the membrane with GOx~ which has previously been activated with an excess of p-benzoquinone, is carried oul. The detailed procedure for this reaction is described in the literature, Sternberg et. al.. Anal. Chem. 1988, 60~ 2781. which is incorporated herein by reference.
B. Physical adsorption of enzvme on cellulose acetate or Nafion followed by crocclinking with glutaraldehvde 1. The sensor is coated with cellulose acetate in exactly the same manner as described above to create membrane. The GOx (270 U/mg) is - . .

WO 91/15993 Pcr/US9l/o26 2 0 ~ O ~ 2 2 physically adsorbed by dropping 5 Ill of GOx solution (40 mg/Ml in 0.1M
phosphate buffered saline) on the in~lira~ing element within cavity 16, and is allowed to dry for 10 minu~es at room temperature. To immobilize the enzyme and form circumferential surface 21, the sensor is exposed to glutaraldehyde vapor generated from 25% glutaraldehyde solution placed at the bottom of an enclosed glass chamber for 12 hours at room temperature. The sensor is then rinsed in de-ioni~d water and dried in air for 2 hours. The croc~1inlcing with glutaraldehyde protects the enzyme from heat degradation, proteolytic enymes and hydrolysis, E.M. Salona. C. Saronio. and S. Garattini (eds), "Insolubilized Enz3nnes." Raven, New York, 1974, ;ncc"~uldted by rcfe.c.,~ herein.
2. Nafion (Perfluorosulfonic acid poly-mer. obtained from E.I. DuPont de Nemours and Co., m~ay~also be used as an ~Itlorn~e fûr cellulose acetale for the inner membrane. After cleaning the sensing portion of the sensor as above, it is electrocoated with Nafion using the method described by Adams e al, Neurosci. Meth VoL 2Z, I987, pp 167-172, illw~la~ by ,er~.~.. ~
herein. One drop of Nafion (5% solution, Aldrich) is placed in a 2 mm loop formed at one end of a copper wire. A DC potential of +3V is applied to the working electrode with respect to the loop for 10 seconds. The sensor is pulled out of the loop before turning off the potential and is dried in air for 2 hours, and the GOx enzyme is applied as described above.
Alternate polymers may be used in lieu of or in combination with cellulose acetate or Nafion for coating of the exposed Pt-Ir wire surface. For example, polyaniline and poly-phenol derivatives can be electroçhemir~lly deposited onto the exposed in~lira~ing electrode surface. Oxidative electro-polymerization of aniline and phenol monomer yields stable and adhesive coating over the exposed wire. These materials moreover have good size selec-tivity which can be utilized to further improve the sensor sele~,livily against electroch.ornir~l inte,~lences in biological envi.ollll.en~. The combination of a size selective coating with a charge selective film (e.g. cellulose acetate) mayreduce the in ~ivo background current and the risk of electroch~mic~l interfer-ence. Electropolymerization ot aniline and phenol is well known, see for exam?le Chsalca ei al. ~lal. Chelll. 19~7, J9, i7~ô-oi, ~nd ivia;;~ L~ e~ ai.
Anal. Chem. l990, 62, 2735-40, both of which are incorporated by reference herein.

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9 20~0~22 Finally; Eastman-Kodak AQ 29-D polymer (poly (ester-sulfonic acid)) has both charge and size selective features, and may be applied ~o the exposed indicating electrode wire in lieu of Nafion. A coating of this type applied to the indicating electrode with a cellulose acetate layer thereover should improveoverall selectivity. Combined coatings made from miAtures of cellulose acetate and the AQ 29-D polymer should also provide advantages in terms of sensor selectivity.
In order to cc F'. e the plepala~ion of the sensor, the whole assembly, including the reference electrode~ is dip coated with 4% polyurethane (Thermedics, SG ~5A) dissolved in 98% tetrahydrofuran (l~ ) and 2%
dimethylformamide (DM~;) to form an outer membrane 24. The polyurethane solution (I0 uL) is held in a wire loop (2 mm i.d.) by surface tension and the sensor is passed through the loop. This leaves a uniform polymer film on the col p'cted sensor 25 to the api!.uy~ate extent depicted in Fig. I. This method l; provides a good control over Ihe arnount of po~rmer which is applied to the sensor. The sensor is dried in air for 6 hours at room temperature and then left in 0.I M i hosph~te buffered saline, pH=7.4 for 72 hours for the various outer membranes to co~ n fully. It is possible to recoat the sensor with polyurethane if the desired linear range of glucose sensitivity is not obtained after the first coating.
EAample 2 - Fii-~. 2 One end of a 10 cm section 26 of Teflon-coated platinum-iridium wire is provided having a 0.I8 mm o.d., a central Pt-Ir wire 28 and a teflon sheath 30 (0.035 mm thi~-lrnl-cs). The left hand end of the wire is stripped to form a ~ - .
cavity 32 as described in EAample I. The right hand end of section 26 is then inserted into a 5 c~ ntimf ~ers long polyethylene tube 34 (0.67 mm o.d., 0.30 mmi.d.). The left hand /_Alre.lli~y of the polyethylene tube is sealed by putting a drop of 4% cellulose acetate solution (in acetone)into the opening. The acetone is allowed to dry while holding the Tefion-coated wire in the middle of the polyethylene tube. This permits ~he formation of a circumferential salt bridge deposit 36 which effectively acts as the terminal part of the reference ciec~rodc, iics in ~ plane tran~ver~e ~ ~h~ ngitu~iin~i axis of the wire ~ anci establishes electrical contact between the reference and sensing electrodes.
The empty annular space between the Teflon-coated wire and the polyethylene tube is then filled under vacuum with 0.1 M phosph ~P buffer, pH = 7.4 con-.
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Wo 91/15993 PCI/US9t/02641 2080~22 -10- ~
- taining 9 glL NaCI. A chloridized silver wire 38 (0.05 mm o.d. 5 cm long prepared as described in Example 1) is introduced inlo Ihe polyethylene tube Erom the righ~ hand end ~hereof and ~his opening is also sealed as described above ~o presen~ a sealing deposit 40. The reterence electrode shows a S potential of -60 + 10 mV (n=6) vs Ag/AgCI (saturated KCL) a~ 370C. The enzyme immobilization and polyurethane deposition steps are then carried out using the p~u._elulcs described in Example 1 to give the inner negatively charged Illclllblane 32a the u h- u~fe~ential in~ir~ing enyme layer 33 and outer permeable luclllb,dne 42 illustrated in Fig. ~. The complete sensor 43 is then ready for calibration and use with electrical ronn~l;o.~ afforded by the aYially ~Yten~ing ends oE the wires 28 38.
The sensors described in the above example are calibrated by dipping into a thermnsrAted cell (aL 370C) con~aining 10 ml of stirred 0.1 M phosphate buffered saline pH = 7.4 and a potential of +600 mV (for hydrogen peroxide leteetir~n) is applied between the work;ing and the . f .ence~ ; e elec-trodes. The background current is allowed to stabilize for 20 minutes. The calibration of the sensor is carried out by adding increasing amounts of glucose~o the stirred buffer. The current is measured at the plateau (steady state response) and is related to the concel ll~tion of the analyte. Following the calibration procedure the sensors are stored in 0.1 M phosphate buffered saline pH = 7.4 at room temperature.
A typical response curve to the glucose addition is shown in Figure 3 for a sensor made in accordance with Fig. 1. As illustrated. the response characteristics of the sensor over the concentration range of interest (0-25 mM) are essentially linear. and are especially so over the range oE 0-15 mM.
The sensor oulput is also essentially independent oE the stirring rate. The in vitro characteristics of the sensor are sullllll .li~ed in the following Table. A
typical storage stability cur~e for the sensor is shown in Figure 4. During the first few days of sensor preparation the polyurethane membrane changes its permeability for glucose as a result of hydrolytic and swelling processes. Ieading to the increased passage of glucose and an increased curren~. After this initialCliGu, L~/Wt;~c;r, ~h~ stability is f YrellPni The sensors of the invention are in use electrically coupled with suitable signal processing e~uipment. and implanted into a desired sl-hcl-~ne-ous site. Glucose and oY~ygen diffusing ~hrough ~he ou~er synthetic polymer .

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membrane are enzymatically catalyzed by the GOX at the in~jca~ing surEace, resulting in production of gluconic acid and hydrogen peroxide. The latter is measured amperometrically, which is a measuremenl ol glucose concentration.

TABLE

In Vitro Characteristics of Fig. 1 Glucose Sensor Parameter Value Residual current (nA~mm~)a 0.7 + 0.2 Sensitivity (nA/mM mm2) 1.~ + O.g Linear Range (upper limit) (mM) 15 ~ 3 Response time (min.), T 90% 3.5 ~ 1 .. ..
Results shown above are expressed as mean + SD for six sensors.
a Residual currents are ll,ca~ul~:d after 1 hour of polarization.

Figs. 5 and 6 illustrate another sensor 44 in accordance with the invention. In this case, the sensor body 46 is received within a stainless steelhollow tubular needle 48.
The sensor body 46 includes an innermost~ Tetlon-coated, pla~inum-iridium wire 50 (90% Pt/10% Ir) having a total O.D. of about 0.2 mm and a cavity 52 formed therein as described in Example 1. The cavity 52 is approxi-mately 1.0 mm in length and is located about 3.0 mm from the tip of the wire 50. A glucose oxidase layer 54 is immobilized within the cavity 52. and com-prises a cellulose acetate polymer layer attached to the surface of the Pt-Ir wire, wilh g,ll~;o~c u)dd~ rl~cciin;cpd thl.Ju~ iuiaraid~hydt: unlu iht:: ceiluiose acetate. This procedure is in acco..lancc with Example 1.B.1. above. The en-tirety of the in~ira~ing electrode is then covered by a membrane 56 oE polyure-thane, again using the method set forth in Example 1.

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... ,. . ~.: - . :-2 0 8 a o 2 2 The sensor body 46 is thereupon inserted into a 25-gauge disposable r - stainless steel hypodermic needle, the latter having an aperture 58 adjacent the l'orward~ sharpened insertion end 60 thereof. The sensor body 46 is installed insuch manner that the glucose oxidase layer 54 comes into registry with the sidewall opening 58, thereby exposing the layer 54 to the biological environ-ment. A silicone rubber plug 62 is installed in the forward end of the needle 48 as shown.
As illustrated in Fig. ~, the wire 50 extends rearwardly out of the end of needle 48, and is adapted to be connected with appropriate instrumenlation ~or I0 measuring glucose conce,.L~atiol1s. In order to seal the rearward end of the sensor 44. a bead 64 of epoxy is applied around the wire 50 and the butt end of the needle 48 and sensor body 46.
The overall sensor 44 is completed by provision of a holder 66 extend-ing transversely of the needle 48. The holder 66 is preferably in the form of a plastic sheet wrapped around the rearward end of the needle 4~ as shown. and secured by means of epoxy or polycyanoacrylate glue. The holder 66 permits ready manipulation and insertion of the sensor 44 even by the patient.
In the use of sensor 44, the reference electrode may be either external-ly applied or implanted. As an external electrode, use may be made of a commercial ele~lu~aldiogram skin electrode described previously may be used.
An external reference electrode should be applied in close plu~dl~ y to the impl~nted sensor for the best measurement results. The holder 66 may also be used to support an external electrode of the type described previously. Inas-much as the holder lies closely adjacent the skin upon implantation, the holder may serve as an ideal platform for the external electrode.
Fig, 7 illustrates an embodiment wherein use is made of an implantable reference electrode. In this case, the needle 48 has an electrodepoci~Pd layer 68 of silver on the external surface thereof, with this layer being ;lnotli7Pd in the presence of chloride ion to create a Ag/AgCI reference electrode. A silver lead wire 70 is conductively affLxed to the rearward end of needle 48 by means of silver epoxy or similar expedient, and the holder 66 is wrapped about this conneciior. aS ;.how.l.
Alternately, the inner wall of the stainless steel needle 48 may be provided with an electrodeposited, anodized silver layer, with conducting gel between this layer and the sensor body 46. A silver lead wire would then be WO 91/1~993 PCl/US91/02641 -13- 2080~22 conducLively secured ~o the inner needle surface. ln this embodiment, electri-cal current flows through the gel between the indicating electrode and the }eferencc f Icc~rod(.
Sensors constructed in accordance with Figs. 5-~, and using either external or implanted reference electrodes, give ~cct-nti~lly the same linear re-sponse as those constructed in accordance with Figs. 1-2.
Actual experience with sensors in acco,dance with the invention has desn-)nctrated that, upon implantation. the cells and capillaries of proximal ~issue are slightly damaged. After four or hve days, however, such tissues I0 lebene.dte around the sensor, forming a collagen layer. Neovascularization has also been observed in the collagen laver. and this phenomenon may partiall,v account for the sensitivity of the sensor. This is i~,di~~ , of o~e~aliol~ of the patien~'s immune system. In any event. the presence of a neovascularized collagen layer adjacent the impl~ntt-d sensor permits passage of oxygen and glucose. In addition, it has been found that in the first hours after implan-tation~ the sensor response is somewhat variable. Over time, however, this variability is de.,-~,ased and the p~.ro,--lance of the implanted sensor ir.~.eases.
This is believed to be due to the s~ r of the tissue around the implant-ed sensor. The end result is that the sensors of the present invention may be successfully implanted and left in place for periods of time heretofore thought a~ al~ e.g., periods of from seven days to three weeks are feasible.
Those skilled in the art will tln~lers~nfl that the sensors oF the inven-tion may require in vivo calibration. This would typically be done by measuring two blood glucose levels by conventional means, and correlating these known values wi~h the ou~pu~ of ~he sensor.
It will thus be seen that the enzymatic sensors in ac~o.dance with the invention exhibit ~.u~..,.~ics h~ ofo.t: difficult to achieve. inrlu-line small, fully implantable size; linearity in response over the concentration ranges of interest;
storage stability; and the abilily to be consistently manufactured without unduerejection rates.

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Claims (13)

1. A sensor adapted for positioning in an environment characterized by the presence of biological molecules which are substrates for or products produced by enzymes in order to determine the presence of said molecules, said sensor comprising:
an elongated flexible body comprising a length of electrically conductive indicating electrode wire covered with an electrically insulative material, there being a portion of said material removed from said electrode wire to define an enzyme-receiving zone;
an enzyme operably immobilized on said zone to present an enzymatic indicating surface;
a reference electrode supported on said body and presenting a section adapted for exposure to said environment, said indicating surface and said reference electrode section being laterally spaced apart along the length of said body and each substantially circumscribing the body, substantially the entireties of said circumscribing indicating surface and said circumscribing reference electrode section being exposed for reaction with said environment when the sensor is placed therein, said indicating electrode wire serving to support said sensor within said environment without the need for a carrier.

2. The sensor of claim 1, said indicating surface being located intermediate the ends of said length of electrode wire with respective segments of said insulating material being on opposite sides of and defining said enzyme receiving zone.

3. The sensor of claim 1, said reference electrode comprising a coil disposed about said body.

4. The sensor of claim 1, said reference section comprising a conductive reference terminal lying in a plane transverse to the longitudinal axis of said body.

5. The sensor of claim 1, including an outer synthetic polymer membrane disposed over said indicating surface and reference electrode section, said membrane being permeable to said biological molecules.

6. The sensor of claim 5, said membrane being formed of polyurethane.

7. The sensor of claim 5, said membrane having a thickness of from about 5 to 10 microns.

8. The sensor of claim 1, including an inner membrane applied to said electrode wire along the length of said enzyme-receiving zone.

9. The sensor of claim 8, said membrane being negatively charged.

10. The sensor of claim 1, said sensor being a glucose sensor.

11. In a method of fabricating an enzymatic sensor including the steps of providing an indicating electrode and operably immobilizing an enzyme on said indicating electrode, the improvement which comprises the step of uniformly coating said electrodes with a synthetic polymer membrane, said coating step including the step of establishing a droplet of polymer solution ofdefined volume, and passing said electrodes through said film to coat the same.

12. The method of claim 11, said film establishing and passing steps comprising providing a wire loop, holding the liquid in said loop by surface tension, to form said film, and passing the electrode through the loop.

13. In an electrochemical sensor adapted to be inserted through the skin of a user and having an elongated, tubular housing presenting a sharpened insertion end and an opposed end with electrochemical sensor means for detecting an in vivo chemical condition carried within the housing and conductor means operably coupled with the sensor means and extending out of said opposed housing end, the improvement which comprises a holder for facilitating manipulation of the housing and insertion thereof through a user's skin, said holder including a transversely extending body secured to said housing adjacent said opposed end and permitting manual grasping and manipulation of the housing, said housing presenting a conductor-clearing opening generally aligned with the longitudinal axis of said tubular housing forpermitting said conductor means to extend out of said tubular housing and past said body.