Microfabricated Potentiometric Electrodes and Their In Vivo

Microfabricated Potentiometric Electrodes and Their In Vivo Applications ErnÖ Lindner Ÿ University of Memphis Richard P. Buck Ÿ University of North Carolina–Chapel Hill

New technologies are yielding mass-produced, miniature, ion-selective electrodes that can be routinely used in biology

D

Potentiometric cells ISEs are chemical sensors that respond to the activity (concentration) of a single ion in the presence of others with the same charge sign. Ion-selective potentiometry

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A N A LY T I C A L C H E M I S T R Y / M A Y 1 , 2 0 0 0

is a routine analytical technique for both in vivo and in vitro analyses, in which the cell voltage is measured as a function of the sample’s solution activity. The general layout of an ISE cell assembly is shown in Figure 1. The electrochemical notation of the cell can be given as

Ag | AgCl | KCl || salt bridge || ε1 ε2 ε3 εj sample inner filling | membrane | | AgCl | Ag solution ε´ ε´´ solution ε4 ε5 εM Cell I

␰

uring the past several years, practical microsensor devices have entered the fields of biology and medicine and are beginning to drive discoveries in cell biology, neurobiology, pharmacology, and tissue engineering. The integration of microfabricated sensors, liquid handling, and biochemical-processing devices with separation systems has led to complete chemical analysis systems on the surfaces of planar devices. However, these “lab-on-achip” devices are envisioned primarily for in vitro, highthroughput screening (1, 2). In this article, we present a small segment of this revolutionary development by discussing the consequences of miniaturizing ion-selective electrodes (ISEs) with respect to their analytical performances and biological applications. Discussed in this article are the most promising trends in potentiometric analysis and thick- and thin-film fabrication processes for mass producing ion-selective and reference electrodes, as well as problems related to the biocompatibility of ISEs and their differences when making in vitro and in vivo measurements. Finally, some examples of in vivo applications will be presented, with a look to the future.

in which phase boundaries are indicated by vertical lines, and two parallel lines mark a liquid–liquid interface. The total potential difference measured between the two terminal electrodes is evidently composed of a considerable number of local potential differences taken from right to left across interfaces.

E = (ε1 + ε2 + ε3 + ε4 + ε5) + εj + (ε´ + ε´´) = Eo + εj + εM

(1)

in which Eo is a constant potential term comprising the potential contributions ε1 to ε5, which arise within the system from the two reference electrodes in the cell; εj is the liquid junction potential; and εM is the membrane potential. We usually seek electrolyte combinations for the reference electrode in which εj is nearly zero. The essential part of an ISE is the ion-sensitive membrane that is commonly placed between two aqueous phases—for example, the sample and the inner electrolyte solution. Ion-selective membranes made from glass, single crystals (e.g., LaF3), pressed pellets of insoluble precipitates (e.g., AgCl, AgBr, AgI), or solvent polymeric membranes (highly plasticized polymeric films, also known as liquid membranes) are widely used in analytical practice. In this article, we will focus our discussions on microfabricated ISEs based on solvent polymeric or liquid membranes. Typically, the membrane potential is divided into three separate potential contributions—the phase boundary potentials at both interfaces (ε´ and ε´´) and the diffusion potential within the ion-selective membrane (3–5). However, the membrane’s internal diffusion potential is zero if there are no concentration gradients within the membrane. The basic equation for the interfacial potentials ε´ and ε´´ is ε=

RT ln kiai z iF a¯i

(2)

in which R is the gas constant, T is the absolute temperature, F is the Faraday constant, z i is the charge number of the primary ion i, ki is the single-ion partition coefficient, and ai and a¯i are the primary ion solution and membrane activities, respectively. The phase boundary potential is a simple function of the sample ion activities only when a ¯ i is not influenced significantly by the sample. Accordingly, for an ideally selective membrane, the membrane potential is directly related to the respective activities in the contacting solutions εM =

RT ln ai(1) z iF ai(2)

E = constant +

RT lnai z iF

(4)

Most of the microfabricated ion sensors are based on solvent polymeric or liquid membranes. Overplasticized poly(vinyl chloride) (PVC) is most frequently used as the membrane matrix. The selectivity of these membranes is determined by the dielectric properties of the plasticizer and the selective, hydrophobic complexing agents, which are neutral or charged carriers (ionophores), within the membrane phase. A comprehensive review of the carrier-based ISEs and bulk optodes was published recently (6, 7). Calibrating ISEs by serial dilution gives a plot of the cell voltage as a function of the logarithm of the primary ion activity, as shown in Figure 2. It is linear over a wide concentration range (generally from 10−1 to 10−6 M) with a 59.2 mV/z i slope at 25 °C. In the linear (Nernstian) response range of the membrane, the composition is constant (8). There are no concentration gradients in the membrane, and the concentration of negatively charged sites is matched with positively charged complexes of the ionophore, which sustains electroneutrality in the membrane bulk, as shown in Figure 2b. Outside the Nernstian response range, the membrane composition changes with the composition of the sample solution. At very low (≤ 10–6 M) or very high (≥10–1 M) sample concentrations, time-dependent concentration gradients develop within the membrane. Interference-induced cross fluxes (Figure 2a) and steady-state salt co-extraction (Figure 2c) are examples of ion transport throughout the membrane and ion leakage into the aqueous solution. Reduced slopes, slow drifts, and higher detection limits accompany iontransport-related changes in the membrane bulk (a¯ i in

Reference electrode

ISE EMF

Ag|AgCl

Ag|AgCl Reference electrolyte

(3)

Inner filling solution

Diaphragm Bridge electrolyte

in which ai(1) and ai(2) are the ion activities in the sample and the inner filling solution, respectively. A Nernstian response is expected in a potentiometric cell when the liquid junction potential can be neglected and the inner filling solution of the electrode is kept constant:

Ion-selective membrane Sample solution

FIGURE 1.Schematic diagram of a potentiometric cell assembly with a membrane electrode and a double junction reference electrode.

M A Y 1 , 2 0 0 0 / A N A LY T I C A L C H E M I S T R Y

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Equation 2) and in the adjoining aqueous solution. In Figure 3, cation interferenceinduced transport in a pH-sensitive membrane is visualized by a novel imaging technique (9–11). The direction of transport is determined by the relative activities in the sample and inner filling solution. The constant release of small amounts of primary ions into the sample controls the activity at the interface (aí or a´í ), which is higher than its nominal value in the bulk. The sensor becomes insensitive to activity changes in the sample bulk, and the theoretical detection limit (DL) cannot be achieved (e.g., blue curve and DLB in Figure 2). However, the surface activity can also drop below the solution bulk value (aí < a i bulk ) when primary ions are transported from the sample toward the inner filling solution. Under such conditions, an apparent “super-Nernstian” response occurs (Figure 2, red curve). Contamination of the adjoining aqueous solution through undesirable transport is especially serious in clinical and in vivo applications for which the sensors are used in ex-

tremely small sample volumes. To approach the theoretically feasible detection limit values, it is essential to eliminate the increased surface concentrations from undesired leaching. The simplest way is to apply flow-through conditions (12) or to use ion buffers in the sample to enforce the nominal bulk activities at the membrane surface (13, 14). Unfortunately, none of these is applicable in vivo. Very recently, the detection limits of ionophore-based liquid membrane electrodes were extended into the picomolar range by eliminating the primary source of the problem, the leaching itself (15). The primary ions are directed through the membrane and into the inner filling solution by using very low primary ion concentrations (10–12 M free ion activity in the presence of EDTA) in the inner filling solutions. Another approach that eliminates the source of leaching is galvanostatic control of transport (12). With fine, programmable, galvanostatic tuning of concentration profiles, it is possible to set the delicate balance between surface concentrations and fluxes in the membrane (16, 17).

Miniaturization and microfabrication

E Ag/AgCl Inner filling solution

DLA

DLC DLB

DLD

log ai

Sample Membrane Inner Sample Membrane Inner Sample Membrane Inner electrolyte electrolyte electrolyte (a) (b) Cl– M+ (c) a´´

N+ abulk M+

M+

M+ R–

a´

N+

R–

+ a´ abulk M X–

a´´

M+ X–

FIGURE 2.Top figure shows schematic calibration curves of a cation-selective liquid membrane electrode, and the bottom figure represents corresponding concentration profiles in the membrane phase. Green shows the detection limit determined by the ionophore properties, blue indicates the detection limit distorted by interference or undesirable ion leaching (Figure 2a), red shows super-Nernstian response that occurred because the surface activity is smaller than the nominal activity in the solution bulk, and black represents the upper limit of detection due to complete salt co-extraction (Figure 2c). (a) Concentration profiles in the ion-selective membrane at very low primary ion concentration; cation interference-induced transport and sample contamination. (aí > aibulk) (b) Concentration profiles in the Nernstian response range. Rust marks the free ionophore concentration distribution. (c) Concentration profiles in the ion-selective membrane at high sample concentrations; transport related to anion interference (co-extraction). M+ and N+ are the primary and interfering cations, respectively; abulk is the primary ion activity in the solution bulk and a´ and a´´ are the ion activities at the membrane surface; (R−) and R− are the intrinsic and intentionally added negative sites, respectively; and ⬍ and ⬍M+denote the free carrier and the carrier–cation complex, respectively.

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Since 1959, when Hinke (18) made early micropipet potentiometric electrodes for piercing single cells and measuring interior ion activities, the miniaturization of conventional diptype, barrel-shaped ISEs have progressed through several stages (19). Simultaneously, progress was made in reducing the sizes of voltammetric redox electrodes made from metals and various forms of carbon. Flat disks, hemispheres, spheres, bands, and other kinds of “ultramicroelectrodes” became fashionable research topics (20). In the early 1970s, flatform, plastic, potentiometric electrodes were perfected by Eastman Kodak’s Ektachem project. By the late 1970s, numerous commercial devices and products for blood ion and blood-gas analysis used flow-through cells with micropotentiometric sensors based on miniaturized dip-style microelectrodes. These are now found in

hospital intensive care facilities and doctors’ offices.

Thin- and thick-film fabrication

(23). The method does not require clean room conditions, and the stamps generated by this method can be reused numerous times without significant loss of performance. Thus, it may become the method of choice for small and medium quantity production runs.

Intensity

Intensity Intensity Intensity

pH 12.7

pH 12.7

pH 12.7

pH 3.9

pH 3.9 pH 3.9 pH 3.9

pH 3.9

Biomedical sensors can be manufactured using many different technologies (21, 22). Microelectronic technology, however, has made it possible to fabricate miniature sensors and even microsensors that can be mass produced. Standard methods for designing and generating microscale patterns on surfaces are generally characterized as thinand thick-film processes. (a) Thin and thick films consist of layers of metals, 40000 insulators, or semiconduc20000 tor materials deposited on insulating surfaces, such as glass, alumina, or polyimide. Thin films can be as thin as a monolayer or have 0 100 200 300 Pixel thicknesses as great as one micrometer. Thick films, (b) on the other hand, have 40000 thicknesses of tens of micrometers. Thin films are 20000 deposited using vacuum evaporation, sputtering, or chemical vapor deposition, whereas thick films are 0 100 200 300 deposited using screen Pixel printing or lamination. (c) The primary limitation 40000 of thin-film microfabrication for sensors is that the 20000 design and capital equipment costs are very high, although per-unit costs may be relatively low. By 0 100 200 300 comparison, the thick-film Pixel technique is relatively inexpensive; however, the qual(d) ity of the deposited film is 40000 lower. Thin-film processes 20000 give patterned films that are highly reproducible (line resolution < 1 µm), whereas thick-film processes yield structures 0 100 200 300 Pixel with greater variability in both in-plane resolution FIGURE 3.(Left) Time-dependent, real color images of the cross section of a ~ 800-µm-thick PVC membrane (~100 µm) and thickness. (ETH 2439/KTFPB/DOS/PVC), during cation interference and (right) the corresponding intensity profiles measVery recently, controlled sagging microcontact print- ured at 450 nm, the absorbence maximum of the deprotonated ionophore. ing has made possible rapid Both sides of the membrane are in contact with a pH 3.9 acetate buffer with 0.1 M Na+ background. (a) At time = 0, the solution on the right side of the membrane is replaced by an equimolar mixture of NaCl and NaOH at pH 12.7, also in 0.1 M in Na+. Times: (b) 9 and inexpensive fabrication min; (c) 2 h, 47 min; and (d) 69 h, 40 min. The protonated ionophore is blue; the deprotonated ionophore is red. Decreased intensities of microstructures, with correspond with increasing concentrations of the deprotonated ionophore. ETH 2439 is a Nile blue-based, pH-sensitive chromoionofeatures down to 1 µm phore, KTFPB is potassium tetrakis-[3,5-bis(trifluoromethyl)phenylborate, and DOS is dioctyl sebacate. (Adapted from Ref. 9.)


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A standard method of fabricating methods. UV light selectively polymerizes the film through patterned thin films in our laboratory is an image on a photographic plate. The unpolymerized porillustrated in Figure 4. First, a thin layer tion of the resist material is washed away. The screen mask of chromium (~200 Å) and gold or stencil is then placed over the substrate, and a thin layer (~2000 Å) is deposited onto a carefully of slurry is forced through the open regions of the mask cleaned substrate surface (Figure 4a) (Figures 5a and 5b). This technique forms a patterned film using vacuum evaporation or sputteron the substrate. The slurry is a liquefied mixture of the ing. The gold serves as a conductor, and the chromium acts powdered film material and a suitable binder. This process as an adhesion layer between the substrate (e.g., Kapton, a is applied using a squeegee whose velocity and force against polyimide film used as a support) and the gold. Next, the the screen are carefully controlled so that a uniform layer of gold layer is spin-coated with a ~1.5-µm-thick positive phomaterial is pressed through the mask onto the substrate. toresist layer (Figure 4b) and selectively exposed to UV light The resultant substrate with the patterned slurry is then through a photomask (Figure 4c) to define the electrode transferred to a furnace and heated to drive away the sites, bonding pads, and connecting traces. The photoresist is binder and sinter the particles into a continuous film. In removed from the exposed areas by immersing the substrate some cases, such as insulating films, the material can actuinto commercial developer. The photoresist remains at areas ally fuse to form a glasslike layer (Figure 5c). where the photomask blocked the UV light (Figure 4d). The major limitation of thick-film devices is that the The exposed areas of gold and chromium are etched away in dimensions of the patterned layers cannot be precisely mainsubsequent steps by submersing the substrate into commertained. The resolution of the screen mask is determined cial etching solutions, leaving the patterned metal layer on by the spacing of the wires in the mesh. The slurry, once the Kapton substrate (Figure 4e). To define the openings for deposited on the substrate, can “run” after it is printed and the electrode sites and bonding areas and insulate the lead while the structure is transferred to the furnace for baking. wires between the two, a thick layer (~30 µm) of photodeThis leads to dimensional changes in the pattern. Neverfinable polyimide is spin-coated over the patterned metal theless, when the dimensional stability is not critical, the layer. It is selectively exposed through a second photomask process is much cheaper than thin-film methods. (Figure 4f). A developer solution is sprayed on the photodeThe precision and resolution of thin-film processing can finable polyimide while it is be approached with thickspinning and until the areas film processing using phounexposed to UV light tosensitive polymer films. Cross section Top Photomask are removed (Figure 4 g). These flexible lamination Additional layers can be foils (e.g., Riston and (a) deposited, creating a mulPyralux) are extensively tilayered structure, as is used in the manufacturing done with thin-film sensors. of printed circuit boards, When doing so, it is imporwith 30-µm lines and (b) tant that etching new film spaces. They have special layers does not affect underadvantages for fabricating lying layers. biomedical sensors and sen(c) The thick-film process is sor systems when combined similar to the thin-film with conventional thin- and Chromium method, and the procedural thick-film processing. LamiGold steps are illustrated in Fignation foils can be used to (d) Photoresist ure 5. The deposition is fabricate and shape sensor UV-irradiated carried out by printing a wells and channels of conphotoresist pattern through a stencil trolled volumes and anchor Polyimide (e) that can be formed on a sensing membranes, protecUV-irradiated polyimide high-density, fine-wire tive layers, gas-permeable mesh or on a thin solid films, and other items to structure with a pattern of the sensor substrate. The (f) holes. In the case of the flexibility of lamination wire screen, a thin, phofoils is another advantage toresist material covers the because ceramic and silicon(g) screen. A photographic patbased devices are too rigid tern is transferred to this for moving living tissue. FIGURE 4.Steps in manufacturing planar electrodes with thin-film material with a procedure The devices shown in similar to standard thin-film processing. Figures 4g and 5c can be

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Cross section (a) Ceramic substrate

Top

Stencil screens

Indirect presensitized film Stainless steel mesh (b) Ag ink

(c) Dielectric ink (glaze)

charge carriers can not pass from one phase into the other (21). To form reversible interfaces on the inner filling solution side of the sensors, hydrogel-filled chambers were integrated into the planar structures. However, incorporating liquid reservoirs with conventional thin-film microfabrication technologies, is technically difficult and expensive. In addition, the overall dimensions of the sensor limit the volumes of wells or chambers built into the structure. Fortunately, with potentiometric sensors, the precision of the patterned layers is often not critical. In such cases, the thick-film process (screen printing or lamination) is more than adequate to form chambers, wells, or channels on the planar sensor platform. Hybrid, thin- and thick-film processing, which are normally not used together, can have unique advantages in sensor fabrication.

Microfabricated reference electrodes FIGURE 5.Fabrication sequence of planar sensors with thick-film screen printing.

considered as general platforms for electrochemical sensor fabrication. To build potentiometric ISEs, the conducting metals in the sensor wells are generally coated electrochemically with Ag|AgCl, and the well is filled with an electrolyte containing a hydrogel, such as polyhydroxyethyl methacrylate, which can be photolithographically patterned. Finally, the appropriate membrane cocktail, dissolved in tetrahydrofuran or cyclohexanone, is dispensed over each sensor well by use of a computer-controlled microdispensor with supports held and transported on x, y, z tables. After complete evaporation of the solvent, generally an additional membrane layer is applied. The total ionselective membrane thickness is ~50 µm. Microfabricated potentiometric sensors are planar versions of the conventional macroelectrodes, reducing in size the typical three-dimensional electrode structures into two-dimensional, multilayer arrangements (Figure 6). The first devices were based on field-effect rather than microFaradic principles (21, 24), using membranes cast on solid surfaces—Si/SiO2 with no intervening internal electrolyte solution. These devices introduced analytical chemists to thin-film microfabrication technologies—photolithographic reduction, thin-film metalization, chemical etching, spin coating, and positive and negative photoresists. However, the reduced size of the sensors in these planar two-dimensional configurations led to problems of potential instability, bad reproducibility, response-time variability, and short lifetimes. Most problems were related to so-called completely blocked interfaces (electrolyte/insulator/metal or electrolyte/insulator/semiconductor), in which the

A microreference electrode has practical relevance when it is built with the same fabrication technology; its stability and lifetime are comparable or better, and its size is at least as small as that of the indicator electrode. Microfabricated electrochemical cells most commonly use Ag|AgCl|KCl reference electrodes. They have excellent stability, as long as the chloride concentration is constant. Exhausting the internal electrolyte from the reference electrode, even at very slow rates, contaminates the sample and influences the potential terms ε2 and εj (Cell I), which leads to drifting potentials and bad reproducibility. Similarly, slow changes in the concentration of the inner electrolyte of the ISE due to water transport through the sensor membrane affects ε´´ and ε4 and, hence, the stability of the cell voltage. To keep the chloride concentration constant, KCl-doped vinyl ester resin (25) was cast over a screen-printed Ag|AgCl reference element, or a KCl-saturated filter paper film was used as a reservoir in a disposable reference electrode (26). Unfortunately, it is extremely difficult to implement reliable, liquid-junction-based, microreference electrodes in chronic, in vivo applications. Therefore, the development of junction-free, all solid-state reference electrodes and solid-contact ISEs of high stability is of great importance (27). Membranes with the same selectivity toward all anions and cations are the ideal candidates for a junctionfree reference electrode. To achieve a membrane-based reference electrode, a cation- and an anion-sensitive membrane were connected in parallel (28), or a membrane was doped with both cation- and anion-exchange sites (i.e., tetraalkylammonium–tetraphenylborate salt) (29). Certain classes of aromatic polyurethanes proved to be especially attractive as a membrane matrix for the fabrication of nonresponsive reference electrodes. Recently, polyion sensors were suggested as junction-free reference electrodes (30). If the valence of


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the polyion is large (z ≥ 60) and no other ions are engaged in ion exchange with the polyion, such electrodes show virtually no response. These novel reference electrodes may have several advantages over the conventional Ag|AgCl|KCl reference electrode if they can achieve the same stability, precision, and accuracy. Redox reference electrode systems are also very promising alternatives, primarily as the inner reference element of the ISE, because their potential is determined by the ox/red ratio, that is, they are not greatly affected by dilution. The Pt|quinone|hydroquinone reference system proved its value as an inner reference electrode for a planar pH sensor (31). The triiodide–iodide solution maintains a redox potential that is constant within 2 mV for periods in excess of two years. Asymmetric, ionic/electronic contact cells, such as

Ag | AgCl | KCl || ε1 ε2 ε3 salt bridge ||

sample | membrane | inert metal solution ε´ ε´´ εM Cell II

involve the transport of ions through one interface and the transport of electrons through the other (32). A permselective membrane for sensing anions or cations, in addition to containing the respective ion-exchange sites, must be loaded with pairs of redox ions—for example, ferri/ferrocyanide in anion-exchanger membranes, or Ru(II/III) complex cations in cation-exchanger membranes. Partially oxidized polymeric films, such as polypyrrole, sandwiched between the membrane and the inert metal, also can mediate charges between ionic (ion-selective membrane) and electronic conductive layers (33).

Water and gas transport effects The Nernstian response of potentiometric sensors does not require extensive motion of free carriers, ion-carrier complexes, or sites. Massive ion transport is expected only in the presence of anion or cation interference. On the other hand, neutral molecules, such as water, O2, or CO2, can be transported easily through the highly plasticized polymeric membranes used for ISE fabrication. The importance of water within the membrane phase was recognized long ago (34–36). It is well known that clear plasticized PVC membranes become cloudy when soaked in water and transparent again when dry. The water content and the water uptake kinetics are influenced by the membrane composition and solution conditions. Both may affect the performance of polymer membranebased planar sensors in several ways. Microfabricated pla-

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nar electrodes need relatively long conditioning periods after first solution contact (“break-in” time), before equilibrium is approached. Long-term potential drifts observed with miniaturized sensors are attributed to slow changes in the equilibrium water content of their membranes (Equation 2) and the underlying hydrogel films, which serve as inner filling solutions. Potential instabilities are observed primarily in samples with large fluctuations in their osmotic pressure. The time needed to bring a microfabricated planar sensor to equilibrium with an aqueous sample is influenced by several parameters. New imaging techniques (11, 34, 37, 38) allow the diffusion of water into dry, plasticized polymeric membranes to be followed. In optimizing integrated, multilayered chemical sensors, the material properties and geometrical features of the membrane and the underlying hydrogel layer are of equal importance. The amount of salt embedded in the hydrogel layer will determine the amount of water necessary to balance the osmotic pressure on the two sides of the membrane—the more salt that is in the hydrogel, the more water will be needed (longer equilibration). Swelling due to the large osmotic pressure in the inner filling solution side eventually may burst or detach the sensing membrane. The water uptake and the well volume should match each other. When the water transported into the inner filling solution side is in excess of the water uptake capacity of the hydrogel (swelling), a thin aqueous layer is formed between the ion-selective membrane and the swollen hydrogel film. Hydrogel layers with little or no salt attract much less water. However, planar sensors with very small volumes of inner filling solution are more sensitive to interferences related to any kinds of material transport through their membranes, which leads to poor potential stability and inadequate reproducibility, similar to coated wire-type electrodes with ion conductor–metal interfaces. Essentially, the required analytical parameters define a practical limit of how far we can go in downsizing microfabricated electrodes with an internal solution (hydrogel) contact. Solid-contact devices can be made much smaller, but their performance generally lags behind those fabricated with aqueous hydrogel inner layers. Other limitations related to the minimal size of these electrodes include a shortage of chemicals in the smaller membranes, increased internal membrane resistance, and reduced contact to the substrate (e.g., Kapton), which leads to poor membrane adhesion. This is immensely important because most device failures are related to poor membrane adhesion, which leads to the formation of lowresistance shunts between the membrane and the substrate. Unfortunately, water content influences polymer membrane adhesion and, thus, the lifetime of multilayered planar sensors (31). Good adhesion is even more critical for tissue measurements so that sensors can withstand cardiovascular motion (39). To increase the adhesive strength, modified PVC membranes with various func-

tional groups, such as carboxylated, aminated, and hydroxylated PVC, are used (31, 39–41). Membranes with no plasticizer (silicone rubber) or decreased amounts of plasticizer (aliphatic polyurethanes) proved to have better adhesive properties in peel-off tests. Fortunately, decreased amounts of plasticizer coincide with reduced host inflammatory response (42). Various gases (e.g., CO2, H2S, SO2, and NO2) and some low molecular weight organic acids (e.g., benzoic and acetic acid) also can pass through solvent polymeric sensor membranes. As they diffuse through the membrane, the chemical composition of the inner electrolyte film, in this particular case the pH, will change (43). This change in pH can be a serious interference, but it also can be exploited for sensing purposes. By selecting appropriate inner solution compositions, pH changes are translated into gas concentrations in microfabricated alternatives to the compound membrane classical gas sensors (44).

In vivo applications

above it. Our sensors were rigid at room temperature and became soft after insertion into the heart tissue. Recording sites are on one side, while reference electrodes are directly behind on the back. This design allows each array to be cooled until it is stiff enough to be inserted into the beating heart. After warm-up, the device is flexible and does not damage heart muscle. The microfabrication steps are summarized in Figures 4 and 5 (39, 40). An advantage of microfabrication is that different kinds of sensors can be integrated onto the same platform using the same or only slightly modified fabrication steps. In a second line of research, we combined potentiometric and amperometric microsensors to follow the accumulation of lactic acid while monitoring H+ and K+ concentrations (45). The measurements were made on the perfused papillary muscle of a rabbit heart, which has been identified as a useful complementary technique for whole-heart, in vivo measurements. The small muscle, ~1 mm in diam and 4mm long, was secured in a custom-designed, transparent plastic chamber with total control of the humidified ambient air. Ischemia was introduced by arresting the flow and decreasing the oxygen tension in the chamber, from 150 mm Hg to 6 mm Hg. A miniaturized, flat-form, conducting, salt-based, amperometric, lactate sensor, using lactate

Our contributions have been aimed at developing thinand thick-film-processing schemes on flexible substrates for fabrication of single sensors and sensor arrays for selective monitoring of ionic and nonionic constituents in biological samples. These are faradic process sensors based on reversible, nonblocked interfaces. They are suitable for in vivo measurements of changing H+, K+, Na+, and Ca2+ concentrations during induced ischemia (localized oxygen deficiency). Cardiovascular physiolo12 mm gists need to obtain and use flexible multielectrode probes with welldefined geometries to describe and understand the electrical activity and ionic distributions in the ischemic heart muscle over time. 4 mm On the basis of our results and prior work, we believe that occlusion of the left anterior descending coronary artery produces discrete regions in which extracellular potasPolyimide sium is elevated to approximately Ion-selective membrane 10 mM, but pH is lowered by no HEMA ISE Reference electrode more than 0.2 pH units. AgCl The array electrode with nine electrode sites and the layer structure Ag of ISE and a reference electrode site Au are shown in Figure 6. The 9-site Cr array is about 1-cm long and 2- to Kapton 5-mm wide on a 125-µm Kapton Solid KCI film platform. A newer version is a double-thick arrangement held together with a memory-shape polyFIGURE 6.At top is the general layout, and below that are cross-sectional views of microfabricated, planar ion-selective and reference electrodes. mer—a material that is rigid below the transition temperature but The Christmas-tree-shaped, nine-site arrays are used with eight sensing sites and one reference site (generally at the tip), or with nine sensing sites and nine reference sites placed back-to-back with the help of a memory-shape polymer. abruptly becomes soft and flexible


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oxidase, operated ideally with essentially zero oxygen partial pressure (46). Microsensors for pH and K+ on flexible Kapton were easily adapted to the rabbit muscle using earlier designs. However, the pH sensor of choice in this application was a special, anodically electrodepositied IrO2 film. These data represent the first simultaneous measurement of extracellular lactate, pH, and pK+ in ventricular tissue during total ischemia. They support the view that pH decreases (due to H+ efflux) and lactic acid transport are coupled. On the other hand, our preliminary observation suggests that K+ transport and lactate may not be stoichiometric.

Biofouling and biocompatibility Ultimately, microfabricated sensors will be used for acute and chronic preparations. Insertion of a biosensor creates an injury and initiates a wound-healing mechanism that is characterized by inflammation. The active surface of the sensor is often negligible when compared with the overall surface area of the implanted device. Accordingly, it is not enough to consider the physical and chemical characteristics of the materials that make up the sensor. The size, geometry, mass, chemical composition, and surface texturing of sensor supports and “accessories” are all critical. To function in vivo, a chemical sensor must meet two requirements. First, it must be biocompatible in the conventional sense—that is, it must minimally perturb the in vivo environment. Second, the environment must minimally perturb the performance of the sensor. The effects of protein adsorption, cellular adhesion, inflammatory reactions, and fibrous capsule formation must be accounted for in devices that are implanted for weeks to months. Phenomenologically, these effects are closely linked. However, the impact of these events on in vivo sensor function can be separated into the deleterious effects caused by membrane biofouling or fibrous capsule formation (47). To prevent membrane biofouling and control the tissue response to an implanted biosensor, complementary strategies can be used, such as modifying surfaces or using special coatings that reduce protein adsorption, applying local drug delivery systems that control tissue responses, optimizing surface topography (roughness, texture, and porosity), and preventing the loss of membrane components into the contacting tissue (42). However, surface modifications of an implanted chemical sensor are more complex than what is required for a simple implant. With amperometric sensors, such as the glucose sensor, the analyte diffusion rate determines the sensitivity, but the selective potentio-

metric response may be completely impaired when ISEs are coated with membrane layers with ion-exchange properties. An adequate analyte diffusion rate is necessary for amperometric sensor sensitivity and short response times.

Perspectives To overcome the medical field’s reticence to implement a new technology, the new should show obvious advantages over the old and justify its cost (48). Modern versions of blood–gas analyzers already use microfabricated, flat-form ion sensors. These state-of-the-art systems also monitor glucose, creatine, creatinine, lactate, galactose, many enzyme activities, hematocrit, and more. In the past few years, the reliability and accuracy of microfabricated, single-use devices have become equivalent to large laboratory analyzers. In parallel, the cost and turn-around time have been significantly reduced, and patient-side testing of critical blood values has become accepted procedure (49).

Microfabricated potentiometric sensors are close to making the move from discrete samples and test technology to continuous, real-time in vivo measurements.

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Microfabrication also offers several unique advantages for in vivo application. Sterilizing wafers and testing a statistically significant fraction of batches before implantation could eliminate the need for elaborate calibrations in a sterile environment. Continued downsizing is expected to lead to a more extensive use of arrays instead of single-element sensors. Sensor arrays using mathematical methods for signal processing could be the solution to various problems that hinder practical systems. For example, multiple data points per sample (per time interval) can provide additional chemical information, eliminate nonidealities, and help discriminate against interferences. Replicate sensors allow corrections for imperfections in fabrication and irreversible reactions on sensor surfaces. A major goal in reducing the size of the implanted sensor is to minimize the injury and inflammatory response. The success is directly correlated to the peculiarities of packaging and the level of integration of control electronics on the microfabricated device. Unfortunately, the lifetime of a potentiometric sensor depends on how long it can resist the loss of critical components. Estimates of residual lifetimes of implanted sensors, based on changes in the membrane composition, are now possible with a chronoamperometric method (50). It has special importance because in vivo calibration of the implanted sensors is not even close to reality, and sensors are tested only pre-

implantation and after removal from the tissue. The use of plasticizer-free membranes with internally tethered ionophores and sites may be the solution for sensors of extended lifetimes in chronic, in vivo applications. In summary, microfabricated potentiometric sensors are close to making the move from discrete samples and test technology to continuous, real-time in vivo measurements. However, a decisive requirement for success will have to be a well-defined need—either established or predicted— in medical practice.

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The work presented here would not be possible without the collaborative work of Ernö Pretsch and his graduate student Titus Zwickl, with support from the U.S.–Hungarian Joint Fund (JFNo:568).

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Ernö Lindner is an associate professor in the joint graduate program in biomedical engineering at the University of Memphis and the University of Tennessee–Memphis. His current research interests include studies of membrane transport phenomena using combined analytical techniques, microfabricated sensors and their in vivo applications, and biocompatibility issues related to chronic electrochemical measurements in tissues. Richard P. Buck is professor of chemistry at the University of North Carolina–Chapel Hill. His research interests include transport of charged species in solids, liquids, bulk, and across interfaces and applications of films and membranes to materials science and sensor technology. Address correspondence to Lindner at the Joint Graduate Program of Biomedical Engineering, University of Memphis and University of Tennessee, Memphis, TN 38152-6582 ([email protected]).


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Microfabricated Potentiometric Electrodes and Their In Vivo

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