Biomedical Sensing Technologies

IEEE SENSORS JOURNAL, VOL. 3, NO. 3, JUNE 2003 251

Emerging Biomedical Sensing Technologies andTheir Applications

Gerard L. Cot, Senior Member, IEEE, Ryszard M. Lec, Member, IEEE, and Michael V. Pishko

AbstractRecent progress in biomedical sensing technologieshas resulted in the development of several novel sensor productsand new applications. Modern biomedical sensors developed withadvanced microfabrication and signal processing techniques arebecoming inexpensive, accurate, and reliable. A broad range ofsensing mechanisms has significantly increased the number ofpossible target measurands that can be detected. The miniatur-ization of classical bulky measurement techniques has led to therealization of complex analytical systems, including such sensorsas the BioChemLab-on-a-Chip. This rapid progress in minia-ture devices and instrumentation development will significantlyimpact the practice of medical care as well as future advancesin the biomedical industry. Currently, electrochemical, optical,and acoustic wave sensing technologies have emerged as someof the most promising biomedical sensor technologies. In thispaper, important features of these technologies, along with newdevelopments and some of the applications, are presented.

I. INTRODUCTION

THE ongoing mergence of the 20th century revolution ininformation technology with the 21st century revolutionin biotechnology poses considerable demand for new sensors,in particular new biomedical sensors. Currently, the recentadvances in the microelectronics industry, the availability ofadvanced microfabrication technologies down to the micro-and nanoscale, and inexpensive signal processing systems havemade the development of a variety of novel biomedical sensorspossible.

A good indication of that demand is the growing use of per-sonal monitoring devices such as glucose sensors for diabeticsor the recently developed sensors for HIV detection. Biomed-ical sensors can also make medical care more personal and tai-lored to the individual needs of a patient. In the near future,a treatment procedure could be adjusted to address a patientsunique metabolism and biological rhythms. For example, thedose of a drug could be correctly determined in order to opti-mize the healing process and minimize its side effects. In addi-tion, biomedical sensors will enable a broad range of medical

Manuscript received May 23, 2001; revised June 10, 2002. The associate ed-itor coordinating the review of this paper and approving it for publication wasProf. Henry Baltes.

G. L. Cot is with the Department of Biomedical Engineering Program,Texas A&M University, College Station, TX 77843-3120 USA (e-mail:[email protected]).

R. M. Lec is with the The School of Biomedical Engineering, Scienceand Health Systems, and the Department of Electrical and ComputerEngineering, Drexel University, Philadelphia, PA 19104 USA (e-mail:[email protected]).

M. V. Pishko is with the Department of Chemical Engineering, Penn-sylvania State University, University Park, PA 16802 USA (e-mail:[email protected]).

Digital Object Identifier 10.1109/JSEN.2003.814656

services at a patients home using a variety of systems that em-ploy a personal computer and the Internet. One may envisiona dedicated home-based analytical diagnostic system interfacedwith a computer that could monitor and store medical data overthe life time of the person. In such a case, specialized applicationsoftware would be capable of recognizing incoming health prob-lems and could notify a person in advance of her or his healthconditions.

A modern biomedical sensor is a device which consists ofa biologically or biophysically-derived sensing element inte-grated with a physical transducer that transforms a measurandinto an output signal. The requirements for any good biomedicalsensor are specificity or the ability to pick out one parameterwithout interference of the other parameters, sensitivity or thecapability to measure small changes in a given measurand,accuracy or closeness to the true measurement, time response,biocompatibility, aging characteristics, size, ruggedness androbustness, and low cost. In addition, the sensor must havecompatibility with the chemical, optical, optoelectronic, orelectronic integrated circuit (IC) technology. The above listedfeatures have been researched comprehensively over the lasttwo decades and critical knowledge has been accumulatedand the challenges have been identified. One may claim thatthe biomedical sensor field has matured enough to be poisedfor commercial success. In this paper, an overview of three ofthe primary biomedical sensor technologies; electro-chemical,optical and acoustic are discussed along with many of thebiomedical applications.

II. MODERN BIOMEDICAL SENSORS

A conceptual model of a biomedical sensor and its impor-tant design elements are shown in Fig. 1. This model presentsa complete biomedical sensor scheme in which, in addition to asensing section of a biomedical sensor, microfluidic, signal pro-cessing and packaging units are included. Simultaneous analysisand design of all these elements are essential for the develop-ment of marketable biomedical sensors.

The principle of operation of such a biomedical sensor canbe inferred by following its sensing path. A measurand is in-troduced to a biomedical sensor using sample delivery systemor by bringing the sensor to the patient, as with implantableor indwelling biomedical sensor probes. Next, the measurandpasses through a preprocessing section, such as semi-permeablemembrane, which performs a initial selective screening of pos-sible interfering factors. After that, the measurand is exposedto the sensing element, a biologically active substance whichis selective to the measurand of interest (i.e., DNA, antibodies,

1530-437X/03$17.00 2003 IEEE

252 IEEE SENSORS JOURNAL, VOL. 3, NO. 3, JUNE 2003

Fig. 1. General diagram of a biomedical sensor system. It should be noted that the biomedical sensor could be an array that would allow for simultaneous detectionof multiple measurands.

enzymes, or cellular components). When a measurand inter-acts with the sensing element, microscopic physical, chemical,and/or biochemical changes are produced. These microscopicchanges cause the macroscopic physical changes in the sensingelement, which are converted by the physical transducer into anoutput electric signal. The electric signal is conditioned, pro-cessed, and displayed. The processing can include such impor-tant sensor features as self-calibration, self-diagnostics, and ad-vanced pattern recognition analyses. All these functional designelements can be enclosed in the sensor package that providesmeasurement integrity to the device.

In recent years significant progress has been made in all areaslisted above. However, the widest researched area has been onthe sensing elements and sensing mechanisms. Below is a shortoverview of the current state in these areas with an emphasison the optical, electrochemical and acoustic biomedical sensingtechnologies.

III. BIOMEDICAL TRANSDUCER TECHNOLOGIESAND APPLICATIONS

Biomedical sensor development has in a large part dependedupon technologies primarily developed for other purposes.The silicon-based microfabrication (IC) and micromechanical(MEMS) techniques have been successfully applied for thefabrication of a large range of miniature electrochemicalbiomedical sensors. Similarly, the progress in optical biomed-ical sensors has been founded on fibers and devices purposedfor fiberoptic communication. Also, acoustic/piezoelectricbiomedical sensors capitalized on the decades-long growth ofRF telecommunication technologies. Piezoelectric elements

utilized in radar, cellular phones, electronic watches, etc.,have been well applied to biomedical sensors. Other types ofbiomedical sensors based on calorimetric, magnetic techniques,etc. have also been effected significantly by modern IC andMEMS techniques.

Currently, electrochemical, optical, and acoustic wave trans-ducer technologies have emerged as some of the most promisingbiomedical sensor technologies. In the following sections, im-portant features of these technologies, some applications andnew trends and developments are presented.

A. Electrochemical Biomedical Sensing TechnologiesElectrochemical biomedical sensors have been the subject of

research for a number of decades, with the vast majority of de-vices consisting of enzymes coupled to an electrode. These de-vices typically can be operated in two modes: potentiometric oramperometric. Amperometric devices are by far the most preva-lent and have seen commercialization for measurands such asglucose and lactic acid. In an amperometric enzyme electrode,an enzyme is immobilized on the surface of an electrode andthe product of the enzymatic reaction detected at the electrodesurface anodically or cathodically. For example, the commer-cially available Yellow Springs Instruments glucose electrodeuses the enzyme glucose oxidase immobilized on a membraneplaced over a platinum electrode. Glucose oxidase catalyzesthe reaction shown at the bottom of the page. One product ofthis reaction, hydrogen peroxide, is oxidized at 700 mV (versusa saturated calomel reference electrode) on the platinum elec-trode surface, producing a current that is directly proportional tothe amount of glucose in sample. Amperometric enzyme elec-trodes for other measurands such as lactate operate in a sim-

COT et al.: EMERGING BIOMEDICAL SENSING TECHNOLOGIES 253

ilar fashion. Amperometric glucose sensors have been by farthe greater commercial success and can be found on most phar-macy shelves in the form of home glucose test meters, such asthe Freestyle meter from Therasense and the One Touch meterfrom Lifescan.

In addition to in vitro enzyme electrodes based on this detec-tion scheme, the FDA has recently approved a subcutaneouslyimplantable glucose sensor [1], [2] based on the same principles.Miniature needle-type amperometric glucose sensors have beencommercialized by MiniMed Corporation (Slymar, CA) forthe short-term monitoring of interstitial fluid glucose (Fig. 2).Despite considerable efforts to develop a membrane and enzymesystem that is minimally affected by the foreign body response,most subcutaneous sensors develop significant drift immedi-ately following implantation. A rapid and progressive loss ofsensitivity can be attributed to protein/cellular fouling of themembrane, enzyme dysfunction due to heavy metal toxicity, andunstable levels of oxygen in the subcutaneous tissues. Accuratereal-time monitoring in the clinical setting requires frequentrecalibration of the in vivo sensor using an external glucose mea-surement reference method [3]. Markwell Medical Corporation(Madison, WI) has developed a miniature implantable enzymebased amperometric sensor with telemetry capable of measuringinterstitial fluid glucose long-term. Unique to this technology is amultilayer protective membrane that promotes angiogenesis andprevents the formation of a diffusion limiting fibrous capsule [4].An iontophoretic transdermal glucose sensor [5][7] developedby Cygnus, Inc. (Redwood City, CA) also used similar glucosedetection chemistry, but uses an electric field to extract interstitialfluid from the skin and uses the sensor to measure glucose in thisfield. Thus, the issue of a foreign body response is removed, butsignificant extraction times (great than 7 min) are required toremove enough interstitial fluid for a measurement.

Medical Research Group Corporation (Slymar, CA) hasdeveloped a glucose sensor for long-term implantation withinthe blood stream [8]. This device is also based on glucose oxidaseimmobilized on a Pt electrode. However, the consumption ofoxygen is measured rather than the production of peroxide.The miniature system looks like a pacemaker generator withtwo flexible intravascular leads. An external patient moduleprovides a data display, visual and audible alarms for hypo- andhyperglycemia, data storage, and telemetry for recalibration. Theprotective membranes are reported to be relatively unaffected byprotein deposition and thrombosis by placing the catheter-basedsensor within the large vein of the chest (vena cava). Oxygensensors are used to measure the decrease in oxygen that occursin proportion to the oxidation of glucose. Venous blood oxygenlevels are measured by a second reference oxygen sensor toprovide enhanced specificity for glucose. Excellent long-termstability has been demonstrated in human testing, requiringinfrequent recalibration. Eventual fouling of the protectivemembranes and enzyme depletion limits the current system tosix months of continuous use. Methods to safely remove theold system and surgically implant a new intravenous system arebeingdeveloped.Anintegratedsensor-controlledinsulindeliverysystem (artificial endocrine pancreas) is being developed byMRG consisting of the vena cava glucose sensor, an implantableinsulin pump and an adaptive computer control algorithm.

Fig. 2. Example of a typical bio-chemical sensing system (minimedcontinuous glucose monitoring system; http://www.minimed.com/files/cgms/patient_pg2.htm).

A number of research programs have sought to minimize theoxygen dependency and improve the signal to noise level ofenzyme electrodes by replacing oxygen in the enzyme catalyzedreaction with an electron acceptor/donor also called a mediator[9]. In the case of glucose oxidase, the reaction becomes thefollowing:

where is the mediator in its oxidized form and is themediator in its reduced form. These mediators are typicallyin the form of organometallic complexes such as ferroceneand its derivatives and Os(imidazole) (bis-bipyridine) .The mediators have lower oxidation potentials than hydrogenperoxide and thus can be operated at lower potentials (200500mV versus saturated calomel). Thus they are less suspectibleto interference from electroactive compounds such as ascorbicacid and uric acid that are pervasive in vivo and in bodily fluids.This type of chemistry has lead to commercial home glucosetest meters such as those manufactured by Abbott Laboratories,Therasense, Roche, and Bayer.

A number of biomedical sensors based upon redox polymerscoupled to oxidoreductases have also been studied. In thesestudies, the polymer served to immobilize the enzyme via for-mation of an insoluble protein/polymer complex, through thephysical entrapment of the enzyme in a polymer film and/orthrough the covalent crosslinking of the enzyme and polymer.These polymers are suitable as mediators for implantable sen-sors because they are less susceptible to mediator leaching fromthe sensor as compared to small molecule mediators. Amper-ometric biomedical sensors based on redox polymer/enzymecomplexes were shown to be miniaturizable and could measuredesired species either intravenously or subcutaneously when im-planted in rats, primates and human volunteers [10][12]. Be-cause of the high current density these sensors exhibit, they canbe calibrated in vivo using a one-point calibration, i.e., thedevice is calibrated from a blood glucose measurement and thesensor current at one point in time.

A significant area recently has been in the development ofbiomedical sensor arrays using nanostructured materials, forboth redundant sensing of a single analyte or multianalytesensing with a single device. Recent research on patterningbiomolecules on surfaces has focused primarily on self-as-sembled monolayers (SAMs) and tethered biomolecules onsurfaces that may potentially form addressable patterned arrays.


Amperometric patterned enzyme electrodes were formed usingSAMs patterned by scanning electrochemical microscopy[13] or via microcontact printing followed by the electrostaticassembly of a multilayer enzyme/redox polymer thin film [14].Willner and colleagues reported amperometric glucose sensorsbased on pyrroloquinoline quinone/glucose dehydrogenaseSAMs that exhibited high current densities [15][19]. Thesesurfaces are easily formed, particularly using alkane thiols andtheir derivatives on gold-coated surfaces. SAMs also permit thesite specific immobilization and orientation of biomoleculeson a surface. However, two-dimensional approaches such asSAMs may limit the number of biomolecule recognition siteson the sensor surface and thus may have low signal levelsand require shielding or other measures to reduce noise. Thestructure of self-assembled molecules on a surface can alsoresult in defects or pinholes in the monolayer and contributeto instability, particularly at applied potentials. The current ad-hesion chemistry used in the fabrication of SAMs also permitsmonolayer formation only on a limited number of surfaces,most commonly gold. In addition to monolayers, photolithog-raphy and other photoinduced patterning chemistries werehighlighted in a few studies, demonstrating the formation ofpatterned biomolecule surfaces and micropatterned polymersfor optical chemical sensing.

Electrode arrays can be fabricated using various methods,however, thin film technology adopted from microelectronicshas been the predominant approach due to ease, quality,reproducibility and low cost of manufacturing. Electrodearrays have been photolithographically microfabricated onthermally oxidized silicon, polyimide, and other insulatingsubstrates. Methods of immobilizing sensing components forglucose and lactate detection onto electrode arrays, includeelectrodeposition of proteins such as glucose oxidase andthe casting of proteins over the array surface. These studiesdemonstrated the potential for fabricating successful sensorarrays, however, there is still a need for a reproducible andsimple way of immobilization of sensing elements which alsoallows for oxygen-free, mediated glucose, and lactate moni-toring. One approach which affords excellent reproducibilityand control over architecture of the deposited species is basedupon multilayer electrostatic assembly [20]. This schemehas been used to deposit oppositely charged redox polymersand enzymes on electrode surfaces. Specifically, ferroceneand osmium based cationic redox polymers were stackedalternately with with anionic oxidoreductases such as pyruvateoxidase (for pyruvate biosensors) and lactate oxidase, both ofwhich as polyanions at neutral pH. Resultant sensors yieldedreproducible current responses and showed no saturation forphysiologically relevant concentrations of analytes such asglucose lactate and pyruvate. Examination of the nanostructureof these thin films suggest that there is significant intercalationbetween the cationic and anionic layers leading to a verycompact and stable nanocomposite structure.

B. Optical Biosensing Technologies

Recent advances in optic and electronic technologies havelead to the research and development of many optic and fiber

Fig. 3. Michelson interferometer using bulk optics. The beam splitter is oftenreplaced with fiber optics.

optic biomedical devices. Indeed, there has been much enthu-siasm as well as a strong effort by newly formed medical de-vice companies, established medical device industry leaders,and universities to investigate optical technologies for a hostof diagnostic and sensing procedures, including both bio-chem-ical and bio-physical based sensors. Examples of some opticalbio-physical sensors include the measurement of body temper-ature, blood velocity and intracranial pressure measurementswhile the optical bio-chemical sensor examples include bloodchemical detection, blood micronutrient monitoring, and cancerdetection. In terms of optical biomedical sensors for monitoringthese parameters remotely into the body, fiberoptic probes canbe used. The utility of fiberoptic probes is that they offer the po-tential for miniaturization, good biocompatability for the visibleand near-infrared wavelengths, fast speed since light is used,and safety, since no electrical connections to the body are re-quired. In this section, we will briefly outline the latest mech-anisms being explored for biomedical optical sensing today in-cluding interferometry, infrared absorption, scattering, lumines-cence and polarimetry.

1) Interferometric Biomedical Sensing: The phenomenonof interference is a method by which physical light interactiontakes place and it depends on the superposition of two or moreindividual waves, typically originating from the same source.There are several variations for producing light interferenceusing both bulk optics and fiber optics but the Michelsoninterferometer is the most common instrument used today(Fig. 3), especially in the Fourier transform infrared (FTIR)machines used for absorption spectroscopy and the relativelynew field of optical coherence tomography [21]. One exampleof a fiber optic based interferometric sensor is the commer-cially available intracranial fluid pressure monitoring systemfor patients with severe head trauma or a condition known ashydrocephalus, which is an increased amount of cerebral spinalfluid in the ventricles and/or subarachnoid spaces of the brain[22]. In addition, the Michelson interferometer is currentlybeing investigated for sensing other biophysical parameterssuch as tissue thickness, particularly for corneal tissue asfeedback for the radial keratectomy procedure, which is laserremoval or shaving of the cornea to correct vision [23], [24].As mentioned, this interferometric approach, when used with


a low coherent light source and scanning or imaging arraytechnology, has been shown to produce morphological datain vivo by a process known as optical coherence tomographic(OCT) imaging [21], [25], [26]. Optical coherence tomographyis a fundamentally new sensing technology that allows visual-ization through tissue at very high resolution. It measures theintensity of backreflected infrared light and allows resolutionsof 1020 m, but with a very limited depth of field, typically12 mm. This approach has been used to investigate severalocular diseases including macular disease, genetic retinaldisease, retinal detachment and retinoschisis, choroidal tumors,optic nerve disorders, and glaucoma [25]. In addition, OCTis useful over approximately the same distance of a biopsyat high resolution and in real time. Consequently, the mostattractive applications for OCT are those where conventionalbiopsies cannot be performed or are ineffective [21]. Thisinterferometric-based imaging approach could also be used incardiology for guiding coronary procedures [27], [28].

2) Absorption-Based Biomedical Sensing: Research in thearea of bio-chemically based optical sensing and, in particular,infrared (IR) absorption-based sensing has been fueled ina large part by the glucose sensing market. However, theoptical monitoring approaches currently clinically availableare the pulse oximeter, IR ear thermometer, hemoglobin andhematocrit meter, and the new IR bilirubin sensor. The pulseoximeter is based on the detection of changes in the strongoptical absorption peaks of oxygenated and de-oxygenatedhemoglobin and are available from vendors such as Criticare,Datex-Ohmeda, Invacare, Novametrix, and Nellcor, to namea few [29], [30]. The ear thermometer is based on receivinginfrared light from the tympanic membrane and is availablefrom vendors such as Becton Dickinson, Braun, Omron, andSafety 1st [31]. Through the work of Groner, Winkelman, etal. [32], and recently commercialized by Cytometrics, Inc.as the Hemoscan 1000 [33], a noninvasive optical approachhas become available which has the potential to monitorhemoglobin and hematocrit as an indicator of iron deficiencyin a clinical or field setting. The approach is based on imagingthe near-infrared absorption spectral peaks of hemoglobin butusing a polarized light input and been referred to as orthogonalpolarization spectral (OPS) imaging. The images are strikingbut the key to this technology becoming truly useful as a fielddevice for blood monitoring is in the software development toprovide quantifiable data. The new bilirubin sensor is a nearinfrared approach pioneered through the early work of Jacqueset al. [34], [35] and commercially developed by scientists atSpectRx Inc.

In general, light absorption in a sample is governed by theBeer-Lambert law in which the transmitted light is dependent onthe wavelength and intensity of the incident light, the path lengthand the absorption coefficient or rather the sum of the multipli-cation of the molar absorptivity times the concentration of all thedifferent components in the measurand [36]. The primary meansfor varying the wavelength of the light include dispersive andnondispersive methods. The optical filter-based nondispersiveNIR systems are more commonly being developed by severalcompanies [37], [38], in particular for glucose sensing, becausethey can be configured inexpensively and generally can have

better throughput if all the light is passed through the sample.However, they have had limited success in producing repeat-able and quantifiable results in vivo. The lack of repeatability ofthe NIR signal in vivo both within and especially between pa-tients is because the signal variations in most instances are notfully understood or accounted for in the system. The primaryknown drawbacks to taking this technology from an in vitro toan in vivo monitoring device include the pathlength variabilityof a pliable tissue, temperature variability of the peripheral sitesuch as the finger or earlobe and the presence of other chemi-cally confounding substances (protein, urea, cholesterol, alco-hols, etc.).

The latest technology in the IR biosensing field is in the de-velopment of clinical and chemical laboratories, in particular atthe Oak Ridge National Lab (ONRL), in which researchers haveminiaturized an infrared microspectrometer to the size of a sugarcube [39]. Carved out of a solid block of plastic, the device mea-sures 1.5 cm on a side and has no moving parts. It can be usedfor blood chemistry analysis, as well as a number of nonmedicalprocesses. The plastic device uses a light source to excite cer-tain types of compounds in gases, liquids and solids. These ex-cited compounds give off infrared light of various wavelengths.The measured emissions wavelengths are fed into a microchip,which determines the concentrations of chemicals in a sample.

3) Biomedical Sensing Using Scattered Light: There arefundamentally two types of optical scattering for diagnosticsand monitoring, elastic and inelastic. The elastic scatteringcan be described using Mie theory (or Rayleigh scatteringfor particles small compared to the wavelength), in whichthe intensity of the scattered radiation can be related to theconcentration, size, and shape of the scattering particles. Thistype of scatter is broadband and not typically specific enoughfor biosensing [36], [40]. However, as a diagnostic screeningtool for cancer detection, measurement of the scatter in thintissues or cells may hold promise [41]. Many of the changes intissue due to cancer are morphologic rather than chemical andthus occur with changes in the size and shape of the cellularand subcellular components. Thus, the changes in elastic lightscatter should occur with morphologic tissue differences. If thewavelength of the elastic light scattering is carefully selectedso as to be outside the major absorption areas due to water andhemoglobin and if the diffusely scattered light is measuredas a function of angle of incidence, there is potential for thisapproach to aid in pathologic diagnosis of disease [36].

One inelastic scattering approach in which the polarizationof the particle is not constant is known as Raman scattering,which is observed when monochromatic (single wavelength) ra-diation is incident upon the media. The shift is associated withtransitions between rotational, vibrational, and electronic levels.As with infrared spectroscopic techniques, Raman spectra canbe utilized to identify molecules since these spectra are charac-teristic of variations in the molecular polarizability and dipolemoments. The Raman signal in general is weak but the tech-nology has advanced, with the replacement of slow photomul-tiplier tubes with faster CCD arrays, as well as the manufac-ture of higher power near infrared laser diodes, to allow forbiomedical sensing. However, this technology is still bulky andexpensive for most biomedical sensor applications and thus cur-


rently in the bench top study phase. The technology has al-lowed researchers to consider the possibility of distinguishingnormal and abnormal tissue types as well as studying a varietyof biological molecules including proteins, enzymes and im-munoglobulins, nucleic acids, nucleoproteins, lipids and biolog-ical membranes, and carbohydrates quantifying blood chemi-cals in near-real time [42][49]. The diagnostic approaches lookfor the presence of different spectral peaks and/or intensity dif-ferences in the peaks due to different chemicals present in, forinstance, cancerous tissue. For monitoring, investigators haveapplied statistical methods such as partial least squares (PLS) toaid in the estimation of biochemical concentrations from Ramanspectra [45], [46].

In addition to normal Raman scattering, biomedical sensorshave been developed using surface-enhanced Raman spec-troscopy (SERS) that allows for orders of magnitude increasein sensitivity. In particular, the ONRL group has developedsurface-enhanced Raman gene (SERG) probes that can locatefree DNA molecules that have hybridized to other DNAs fixedon a surface [39]. This group uses a hybridization method inwhich DNA is transferred from the nylon membrane to a glassstrip coated with tiny silver spheres. The dye labels attachedto the DNA bases have unique Raman infrared spectra, but thenormally weak Raman lines are greatly enhanced by the presenceof the silver spheres. This enhancement allows DNA bases to bedetected at sufficient sensitivity to be useful for DNA sequencingstudies. Using a silver colloid, SERS has also been used to detectelevated levels of amino acids, in particular glutamate, in cerebralspinal fluid after head trauma [50]. These elevated levels couldcause further damage and thus need to be quantifiably measuredin order to provide for pharmaceutical intervention.

4) Use of the Luminescence Property of Light for BiomedicalSensing: Luminescence is the absorption of photons of electro-magnetic radiation (light) at one wavelength and re-emission ofphotons at another wavelength. The luminscent effect can bereferred to as fluorescence or phosphorescence. Fluorescenceis luminescence that has energy transitions that do not involvea change in electron spin and therefore the re-emission occursmuch faster. Consequently, fluorescence occurs only during ex-citation while phosphorescence can continue after excitation.

The measurement of fluorescence has been used for both di-agnostic and monitoring purposes. Obtaining diagnostic infor-mation, in particular with respect to cancer diagnosis or thetotal plaque in arteries, has been attempted using the intrinsicfluorescence of tissue [51], [52]. The intrinsic fluorescence isdue to the naturally occurring proteins, nucleic acids, and nu-cleotide coenzymes while extrinsic fluorescence is induced byseveral mechanisms. For instance, antibody-based extrinsic flu-orescent biomedical sensors have been developed for numberof biomedical applications including cancer diagnosis [53]. Infact, such fluorescent biomedical sensor probes can be made atthe nanometer scale in which a laser beam (blue) that penetratesa living cell is used with monoclonal antibodies that recognizeand bind to benzo(a)pyrene tetrol (BPT), indicating that the cellhas been exposed to a cancer-causing substance [54]. Extrinsicfluorescence has also been investigated to monitor such measur-ands as glucose, intracellular calcium, proteins, and nucleotidecoenzymes [36]. Unlike the use of fluorescence in dilute solu-

tions, the intrinsic or autofluorescence of tissue as well as thescattering and absorption of the tissue act as a noise source forthe extrinsic approach.

Most extrinsic fluorescent biomedical sensors are based oneither the measurement of intensity or lifetime, in which thelifetime can be measured in the time or frequency domain[55][58]. Bench-top systems typically are bulky and includedual monochromators (grating based wavelength separationdevices) are used with either a photomuliplier tube as thedetector or a CCD array detector. However, once the optimalconfiguration for a particular biomedical application, suchas cervical cancer detection or glucose sensing, has beeninvestigated using the bench-top machine, an intensity or phaselifetime measurement system can be designed with a simpler,more robust, configuration. Such a system can be designedwith wavelength specific filters instead of monochromators andmade to work at two or more discrete wavelengths. In addition,optical fibers can be used for delivery and collection of thelight to the remote area [58]. Since the excitation wavelengthand fluorescent emission wavelengths are different, the samefiber or fibers can be used to both deliver and collect the light.

As with many of the optical approaches discussed, a numberof novel fluorescence-based techniques seem to have evolvedfrom the glucose sensing application. Unlike many of the other,noninvasive, optical approaches being investigated, because thefluorescent glucose sensing approaches have to be in contactwith the sample, they have the advantage of being highly spe-cific to the glucose. Those approaches that seem to have demon-strated the most promise generally fall into two categories: theglucose-oxidase based sensors and the affinity-binding sensors[43]. In the first category the sensors use the enzyme catalyzedoxidation of glucose by glucose-oxidase (GOX), similar to theelectrochemical sensors but in this case they generate an opti-cally detectable glucose-dependent signal. Several methods foroptically detecting the products of this reaction and hence theconcentration of glucose driving the reaction have been devised[59], [60]. The primary drawback to GOX based sensors is thattheir response depends not only on glucose concentration but onlocal oxygen tension as well [43].

The affinity-based sensors do not depend on local oxygen;however, many of the earlier affinity-binding techniques wereinvestigated for short term use since they required indwellingprobes [57], [58], [61][63]. In more recent work, these andother investigators have exploited the phenomenon of fluores-cence resonance energy transfer (FRET) whereby an acceptor inclose proximity to a fluorescent donor can induce fluorescencequenching in the latter. Most recently, Russell et al. has reportedthe use of poly(ethylene glycol) or PEG particles to encapsulatethe FRET assay [55]. In the work of Russell, it was reported thatit is possible to create a microparticle-based fluorescent glucoseassay system potentially suitable for subcutaneous implantation.

In addition to the above application, high sensitivity, shortcollection time requirements, lack of sample contact, andcapability of scanning large areas/volume render fluorescencemethods an attractive alternative for microbial detection. Onefluorescence-based approach uses the intrinsic fluorescenceof tryptophan, other amino acids, and DNA which are excitedand fluoresce in the UV (200300 nm excitation, 300400 nm


Fig. 4. Schematic diagram of a biochip system (adapted from the virtual poster presentation of ORNL [http://www.ornl.gov/virtual/biomedical sensors/]).

emission). Observation of these markers is a sensitive indicatorof the presence of biological materials. However, all biologicalmaterials contain these building blocks and, in addition, par-ticles, such as dust, pollen, smoke, etc., preferentially scatterUV/blue radiation (Raleigh scattering). These interferencessignificantly reduce the utility of these markers. However, avariety of microbial cell components exhibit intrinsic fluo-rescence: NAD[P]H, tryptophan, tyrosine, DNA, lumazines,pterins, flavoproteins, and other secondary metabolites [64]. Ofthese, NAD[P]H (360 nm excitation, 450/475 nm emission) hasbeen the most extensively studied. Other DPNH metaboliteshave excitation and emission energies near that of NAD[P]Hand these metabolic signals can be obtained by integrationof the 425500-nm region fluorescence. This fluorescence isdirectly proportional to the concentration of the metabolite,e.g., NAD[P]H and related to the number of live (metab-olizing) cells. Other components, such as flavoproteins andcytochromes, have some absorption in this region, but theirconcentrations are factors of 10100 less in living cells. In fact,there are no interferences of biological origin in this region andthis fact has made redox fluorimetry based on NAD[P]H/DPNHa vital tool in the investigation of cellular metabolism andtissue oxygenation [65]. One method and prototype device forthe detection of microbial life on surfaces, such as foods, glass,plastics, cloth, stainless steel, etc., developed by Estes et al.,has resulted in a sensitivity of in environmentalconditions [66].

Last, the latest research in luminescent sensing is a newgeneration of reagents that report on specific molecular eventsin living cells, called fluorescent protein biomedical sensors orrather bioluminescent chips. These biomedical sensors haveevolved from in vitro fluorescence spectroscopy and fluorescentanalogue cytochemistry. The various probe designs measure themolecular dynamics of macromolecules, metabolites, and ionsin single cells emerge from the integrative use of contemporarysynthetic organic chemistry, biochemistry, and molecularbiology [67]. For instance, a simple device consisting of acoating of bioluminescent bacteria on top of a light-sensitiveintegrated circuit can be used to emit light in direct correlationto an measurand concentration [68]. The overall biochip designcan be represented as shown Fig. 4, as adapted from the virtualposter presentation of ORNL [69].

5) Biomedical Sensing Using Light Polarization Proper-ties: Two of the emerging applications of polarized lightare for biochemical quantification such as glucose and tissuecharacterization, in particular, to aid in cancer identification[70][73].

The concept behind these devices for measurand quantifica-tion is that the amount of rotation of polarized light by an opti-cally active substance depends on the thickness of the layer tra-versed by the light, the wavelength of the light used for the mea-surement, the temperature, the pH of the solvent, and the con-centration of the optically active material [43]. For polarimetryto be used as a noninvasive technique, for instance in blood glu-


cose monitoring, the signal must be able to pass from the source,through the body and to a detector without total depolarizationof the beam. Since the skin possesses high scattering coeffi-cients, which causes depolarization of the light, maintaining po-larization information in a beam passing through a thick piece oftissue (i.e., 1 cm), which includes skin, would be very difficult,if not impossible, although at least one company is attemptingthis approach [74]. Tissue thickness of less than 4 mm that in-clude skin have been tried [75] and may potentially be used butthe polarimetric sensing device must be able to measure millide-gree rotations in the presence of greater than 95% depolarizationof the light due to scattering from the tissue. As an alternative totransmitting light through the skin, the aqueous humor of the eyehas been investigated as a site for detection of in vivo glucoseconcentrations since this sensing site is a clear biological opticalmedia [70], [71]. It is also known that glucose concentration ofthe aqueous humor of the eye correlates well with blood glu-cose levels, with a minor time delay (on the order of minutes),in rabbit models [71]. The eye as a sensing site, however, is notwithout its share of potential problems. For instance, potentialproblems with using the eye include corneal birefringence andeye motion artifact [70]. In most tissues, including the eye, thechange in rotation due to other chiral molecules such as pro-teins needs to be accommodated in any final instrument. In ad-dition, most other tissues also have a birefringence associatedwith them that would need to be accounted for in a final polari-metric glucose sensor.

It is the birefringence and retardation of the polarized light, aswellaspolarizedscatteringof thecellsand tissue, that is thesignalrather than the noise when using polarized light for tissue charac-terization. It has been shown that scanning laser polarimetry pro-vides statistically significant higher retardation for normal eyesin certain regions over eyes with glaucoma. In addition to retar-dation in the retinal nerve fiber layer, images generated from thescatteringofvariousformsofpolarizedlighthavealsobeenshownto be able differentiate between cancer versus normal fibroblastcells in vitro [76]. Using a simplified, cross-polarized system, thepotential of polarized light to not only distinguish between pig-mented nevus and a freckle but also to retrospectively identify thetrueborderaroundapatientwithasclerosingbasalcell carcinomahas been tested in vivo [73], [77].

C. Acoustic Biomedical Sensors

Acoustic waves have been used to study physico-chemicalproperties of gases, liquids, and solids for decades [78], [79].In recent years, there has been an increase in efforts to utilizeacoustic waves for the development of biomedical sensors[80][83]. Acoustic wave transducer technology providesa wide range of devices that are sensitive, accurate, small,portable, robust, and have excellent aging characteristics. Inaddition, such transducers can be produced using standardphotolithography and hence are inexpensive. Acoustic wavescan be generated and received by a variety of means includingpiezoelectric, magnetostrictive, optical, and thermal techniques[84]. The piezoelectric transduction effect almost exclusivelyhas been utilized for generation and reception acoustic wavesin sensor applications.

Acoustic biomedical sensors are typically designed to operatein a resonant type sensor configuration implemented as an os-cillator. In this case, the output sensor signal is the resonant fre-quency shift, which is a function of the magnitude of the mea-surand. This is an important feature, since one can measure fre-quency relatively easily. In addition, the frequency can be con-sidered as a quasidigital signal, which facilitates the subsequentsignal processing operations. Sensors configured as an oscil-lator can be equipped with an antenna for remote sensing andcontrol. Several recent oscillator circuits include an option formeasuring sensor losses which significantly expands measure-ment capabilities of portable acoustic sensors [80], [85]. An-other advantageous feature of using piezoelectric materials isthat the same electro-mechanical transduction mechanism canbe used not only for sensing but also for actuation. This propertyis essential for the design of piezoelectric surgical microcutters[86] or liquid microflow systems [87]. Therefore, it is possibleto develop smart biostructures in which both sensing and actu-ating are realized within the same piezoelectric technology plat-form. Recently, a piezoelectrically-based Chem-Lab-on-a Chipcapable of detecting a variety of hazardous chemicals has beenreported [88]. In summary, the piezoelectric sensing platformoffers a very versatile technology base for the development ofsensors, actuators and smart structures.

1) Acoustic Waves, Piezoelectric Transducers and AcousticSensing Mechanisms: There are different types of acousticwavesthatcanbeusedforbiomedicalsensing.Knowledgeoftheirproperties is important for the selection of the optimal acousticwave for a given measurand. Acoustic waves can be consideredas a source of distributed force acting on a medium. The resultantdeformation of a medium can be compressional or shear, orconsists of a combination of both. The type of the deformationaccompanying the wave is important because it determinesresultantacousticsensingprocesses.Compressionaldeformationis associated with the structural relaxational processes in themedium, while shear is coupled with medium viscoelasticproperties and therefore they are sensitive to different molecularprocesses [78], [79]. Compressional deformations are easilytransmitted through any gaseous, liquid or solid media. Sheardeformation, on the other hand, propagates only through solidsand penetrates only into liquids and gases. This last featureis very favorable because it makes shear waves sensitive toa numerous interfacial phenomena and this configuration isutilized in most acoustic wave biomedical sensors. The wavepenetration depth, which depends on the frequency of the waveand the density and viscoelastic properties of the medium,ranges from microns to nanometers. Therefore, required samplevolumes for sensing are small and the sensitivity of the sensoris high. In addition to mechanical phenomena, acoustic wavesensors can also sense electrical properties of a medium. Electricfield probing of a medium is either generated by acoustic wavedisplacement (via the piezoelectric effect) or is provided by asensor electrode structure. As a result, acoustic wave sensingmechanisms are very broad and acoustic sensors are capableof measuring the changes in mass/density, elastic modulus,viscosity, electrical conductivity, and dielectric constant.

There are several types of waves that can be excited bypiezoelectric transducers and they can be classified as bulk and


Fig. 5. Classification of acoustic waves (bold-marked acoustic waves already have been used in biomedical sensors).

surface generated waves. The bulk generated waves are usuallyexcited by metalized bulk piezoelectric elements such as disksor rods, whereas the surface generated acoustic waves areexcited by an interdigital system of metallic electrodes (IDT)placed on the surface of piezoelectric materials. The electrodesare used to connect the transducer with electronic circuitryfor the excitation and/or reception of acoustic waves. Two ofthe most common configurations of piezoelectric transducersinclude a thin metalized disk used for the excitation of bulkwaves which excite thickness shear modes (TSM) and aninterdigital transducer (IDT), which is applied for the excitationof a surface Rayleigh wave (SRW), a surface transverse wave(STW), a shear-horizontal acoustic plate mode (SH-APM), anda flexural plate wave (FPW). The STWs form a large family ofvarious waves and include shear horizontal SAW (SH-SAW),surface skimming bulk wave (SSBW), and Love wave modes[83], [90]. In Fig. 5, a general acoustic wave classificationdiagram is presented. Knowledge of the properties of acousticwaves is important for the optimal design of a biomedicalsensor [80], [89], [91]. Specific properties of these waves, suchas the type of accompanying mechanical displacement, spatialdistribution of mechanical and electrical fields, susceptibilityto spurious mode coupling and the sensitivity of the wave toambient conditions such as temperature, pressure, etc., dependon the cut of piezoelectric materials from which transducers aremade. There are many available materials for manufacturingpiezoelectric transducers including crystals, composites, andhybrid structures, which provide a wide range of possiblesensing material design options. The most often used materialis quartz, which is chemically inert, has superior mechanicalproperties and is temperature compensated [80], [81]. AT-quartz cut orientation is routinely utilized for fabricationof TSM-based resonator biomedical sensors. In Fig. 6, aschematic presentation of a distribution of shear mechanicaldisplacement generated by a disk-shaped AT-cut transducerimmersed in water is given when the transducer operates at thefundamental (a) and harmonic frequencies (b). As a generalrule, acoustic transducers can slice a biological interface at

Fig. 6. Conceptual model for a TSM transducer exposed on one side to water.(Top) Sensor operating at the fundamental frequency. (Bottom) Sensor operatingat the fundamental frequency and higher harmonics.

different depths [Fig. 6(b)], hence providing important spatialinformation. For SAW based devices various cuts of ,

and crystals are used [80], [83]. Here, commonlya 36Y rotated [82], ZX [83] and ST-90[86] materials are utilized. In the case of the APF sensors asilicon-ZnO hybrid structure is employed [80].

In a typical biomedical sensor, piezoelectric transducers areintegrated with biological sensing films in order to obtain arequired specificity (Fig. 1). Many biological materials suchas proteins (enzymes, antibodies, receptors), organelles, cells,and tissue (microorganizms, animal and plant cells, and tissues)have been used as molecular recognition elements in piezo-electric biomedical sensors. An important issue impacting the


Fig. 7. Typical acoustic sensing process.

sensor performance is a proper attachment of a biological film tothe sensor surface. In addition to classical immobilization tech-niques [90], several novel bio-deposition techniques have beenrecently developed and are successfully used with piezoelectricsensors. Examples include SAMs [92], molecularly imprintedpolymers (MIPS) techniques for immobilization of antibodies[93], and soft lithography patterning for proteins and cells [94].

A typical acoustic sensing process is schematically presentedin Fig. 7 and the corresponding measurement systems aredepicted in Fig. 8. When a measurand interacts with the sensingelement, microscopic physical, chemical and/or biochemicalchanges are produced. These microscopic changes cause themacroscopic acoustical/mechanical or electrical changes in thesensing element. Specifically, the density, viscosity, elasticity,electric conductivity, or dielectric constant of the sensing ele-ment undergo changes, which in turn modify the acoustic fieldquantities of the acoustic wave transducer. The acoustic wavetransducer, which consists of a piezoelectric element with anarray of metal electrodes, acts as a converter which transductsmeasurands into an output electric signal. For example, inthe acoustic immunosensors, the antibodies are immobilizedin the form of a thin film at the surface of the acoustic wavetransducer. When the target antigen is introduced into thesensor environment, the elasticity, density, and viscosity of thefilm vary and these variations modify the acoustic parametersof the sensor, which finally leads to the changes in the outputsensor signal. Molecular interpretation of the sensor responsecould be very simple or very complicated. When the sensingfilm is rigid and acoustically thin, then the sensor response canbe related directly to mass accumulation of the measurand atthe interface [80], [96]. Examples include gas absorption by

thin rigid polymer films [83] and thin metal oxide films [97], orsome immunological reactions [98]. However, when the filmis acoustically thick then, in addition to mass effects, the filmviscous and elastic properties make significant contributionsto the sensor response and the relationship between the sensorresponse and molecular events could be very complex [80],[99], [100]. Also, other factors such as the type of the boundaryconditions or the topography of a sensor surface come intoimportance and influence sensor response, as well [101], [102].Therefore, different biological systems and measurementconditions should be carefully studied in order to select or todevelop the correct sensor model representation. Modeling ofthe sensor response is an area of ongoing, very active research,and though significant progress has been made in the last sev-eral years, many issues still need to be addressed [103], [104].Among them are the molecular nature of the electro-bio-me-chanical boundary conditions, interfacial viscoelasticity, andthe development of dedicated computational techniques.

Acoustic biomedical sensors, from an electronic measurementpointofview,canbe referred toaselectrical radio-frequency(RF)components such as resonators, filters, or delay lines. Therefore,standard laboratory microwave measurement techniques basedon network analyzers, vector voltmeters, and impedance metersare routinely used for the characterization of biomedical sensors(Fig. 8). However, the resolution of these techniques is limitedand they are not applicable for real-life applications because ofthe cost and size. Therefore, in recent years, a lot of effort hasbeen placed on the development of small portable biomedicalsensor measurement systems. Several designs allowing simulta-neous measurement of the transducer frequency and dissipationchange have been proposed that are based on the frequencyand amplitude measurements (oscillators) [85], the decay timeand amplitude [87] or their combinations [91], but the overallperformance of these systems is only satisfactory [80]. Finally, avery difficult and rather neglected issue is that of an enclosure fora sensor. The separation of a fluid-based biological environmentfrom the electronic circuitry poses a very difficult task. Thoughseveral ingenious mechanical measurement cell designs havebeen devised [104] and various electrical passivating techniquesutilizing sputtered thin layers [96] and deposited polymerfilms [97] have been successfully tested, the complete designthat is commercially viable with an appropriate fluidic deliverysystem still needs to be developed.

2) New Acoustic Sensing Devices and Applications: Anovel monolithic piezoelectric sensor (MPS) has been recentlypresented for the detection of chemical and biochemical mea-surands [105]. This new sensor overcomes several deficienciesassociated with the typical two electrode system of a bulk TSM(QCM) sensor, while still employing a well-characterized,temperature-stable thickness-shear mode (TSM) response. Thisnew three-electrode structure, depicted in Fig. 9, is applicableto both gaseous and liquid phase measurements; however, theprincipal benefit of the MPS is in liquid phase measurements.In these applications, it offers the ability to operate simple,yet stable, oscillator circuits in relatively viscous media. Thisnovel MPS structure should accelerate the commercializationof piezoelectric sensor technology, particularly in such areas asbiomedical, biochemical and environmental testing.


Fig. 8. Laboratory and portable electronic measurement systems utilized withacoustic biomedical sensors.

As well as the acoustic transducer technology, novel in-terfaces are also being developed. Biological recognitionelements require appropriate immobilization on the surface of atransducer in order to maintain their bioactivity and specificityof biointeraction. Several immobilization techniques, surfacemodification methods, and a variety of novel intermediate ma-terials supporting the attachment have been studied extensively[106]. In particular, several polymer-based interfaces have beenproposed [107], [108]. One technique is a new very promisingpolymer nanofiber structure for cell immobilization [109].Nanofiber architecture closely resembles a natural collagenstructure which supports and promotes growth and proliferationof tissue cells.

In addition to the aforementioned novel acoustic sensordevice and the artificial biomedical sensor interfaces, therehave been several new applications for acoustic sensors. Theseinclude antibody-based, nucleic acid-based, enzyme-based,and cell-based acoustic biomedical sensors. Antibody-basedor immunosensors have been the most explored and the mostadvanced class of acoustic biomedical sensors. The interestresults from the fact that antibodies can be obtained againstalmost any substance, antibody immobilization techniques arenumerous and well developed and the pertinent acoustic sensingmechanism is simple and mostly determined by mass accumu-lation [110]. The detection process of antigens does not requirelabeling, which is the obvious advantage. However, in a casewhen nonspecific binding may interfere with detection, specialcare needs to be taken to eliminate it [111]. Other advantages

(a)

(b)Fig. 9. Physical structure of the MPS. (a) The two active electrodes (uppersurface) each have area, A and are seperated by a gap width g. (b) The MPSequivalent circuit consists of two TSM equivalent circuits with mechanicalcoupling (modeled as mutual inductance) and a parasitic capacitance. (c)Experimental amplitude (solid) and phase (square) response of the prototypeMPS with various viscous water-glycerol mixture loading [105].

of piezoelectric immunosensors are that they are reusable withno noticeable degradation in performance [112]. In addition,piezoelectric immunosensors can operate in optically opaquemedia. Furthermore, piezoelectric sensors are inexpensive,easy-to-use, and feature rapid response; hence, they may allowfor wide screenings and the development of effective preventivestrategies for a broad range of diseases. Such sensors are usedto determine the concentration of immunoglobulin IgM andC-reactive protein [112]; human cells such as T-lymphocytes[113], erythrocytes [114], herbicides in drinking water [115];bacteria such as E.coli [116], Stalphylococcus aureus [117];drugs (methamphetamine) in human urine [118]; and virusessuch as human herpes [119], hepatitis [120], African swinefever [121], and HIV [122]. In particular, piezoelectric sensorsfacilitated the detection of antibodies that are specific againstHIV within a few minutes in human serum sample. It is worthnoting that the selectivity and sensitivity was on a par with thatof a licensed HIV ELISA and a typical HIV sensor responseis given in Fig. 10 [122].

Exploiting advances in genetic sequence information tech-nology and utilizing an inherent strong affinity interactionbetween complementary nucleic acid strands, nucleic acid,or DNA sensors are emerging as a very important class ofbiomedical sensors [123], [124]. Unlike immunosensors,which are susceptible to nonspecific binding, DNA sensorsprovide higher selectivity and reliability. Similar to acousticimmunosensors, no label is required for the detection of DNA.However, much more complex immobilization procedureschallenge development of DNA sensors. DNA biomedicalsensors typically employ an immobilized single-stranded DNAmolecule for hybridizing with the complementary strand ina given sample. Recently, a TSM quartz sensor was usedto monitor hybridization and affinity of several synthesizedantisense oligonucleotides [125]. Antisense oligonucleotidesare potential therapeutic agents used against cancer and viral


(a)

(b)Fig. 10. (a) Resonance frequency change versus the time of a TSM HIVimmunosensor and (b) comparison of the results measured with the TSM HIVsensor and with licensed ELISA [122].

diseases. A combination of a piezoelectric biomedical sensorwith PCR-amplified real Aeromonas bacteria samples alloweddetection of hybridization using only 205 bp fragments ofextracted DNA [126]. Genetic defects were identified with aTSM biomedical sensor by measuring a single mismatch in a15mer single-strand target of the p53 suppresser gene [127].Other work reports a successful study on denaturation of DNA[128], the detection of DNA polymorphisms [129] and thereal-time monitoring of enzymatic cleavage of nucleic acids[130]. In comparison to traditional DNA analysis methods[123], acoustic sensors offer a rapid and marker-free detectionof nucleic acid hybridization. Acoustic DNA sensors shouldfind a broad application in medical area for fast and inexpensivescreening of variety of diseases, in the pharmaceutical industryduring drug development process and also address manysensing needs in monitoring environmental conditions.

Enzymes provide another important interface capable ofhighly specific reactions with a variety of biological substances.Though enzymes are biocatalysts, they participate in biologicaltransduction processes and have been extensively used forbiomedical sensor development. Enzymes catalyze bioreac-tions at a very high reaction rate, have well-characterizedmechanisms of action, are easily immobilized, are availablefor broad range of applications and are inexpensive [131].Several AT-cut TSM quartz glucose sensors using immobilizedhexokinaze [132] and glucose oxidase [133] have been studied.In the last case, glucose levels were measured in situ and thesensor responded in 80100 s with a linear relation to glucose

in the concentration range between 30200 uM. Chang andShih [134] developed a fullerene-cryptan-coated piezoelectricmembrane glucose enzyme sensor by measuring gluconic acid,a product of glucose oxidase in glucose aqueous solutions.They found that the interference from various common speciesfound in the human blood was negligible. Piezoelectric glucosemeters are in a position to pose an alternative solution to wellknown electrochemical portable glucometers.

Last, acoustic biomedical sensors incorporating living cellsare capable of delivering functional information in contrastto previously discussed protein-based sensors, which provideanalytical data. Functional information, i.e., information aboutthe physiological effect of a measurand on a living system, isoften desired in many important applications in pharmacology,toxicology, cell biology, and environmental measurements[135]. Important properties of cultured animal cells such asattachment, proliferation and cell-substrate interaction underdifferent conditions were successfully monitored using TSMsensors. In particular, extracellular matrix deposition, theintegrity of the actin cytoskeleton, cell-substrate separationdistance, and mechanical properties of the narrow cleft be-tween cell and substrate were determined [136]. Steinem et al.[137] utilized acoustic sensors for analysis of the interactionsganglioside-lectin and ganglioside-toxin (cholera, tetanus, per-tusis). Because a cell-based sensing process is physiologicallyrelevant to natural cellular machinery, these sensor types willexperience growing significance in the near future.

IV. CONCLUSIONAbout $20 billion per year are spent for analytical testing

worldwide. Specialized laboratories located away from apatient, doctor, or hospital perform nearly all of the testingcausing significant time delays in reporting results. Modernbiomedical sensors developed with advanced microfabricationand signal processing techniques are becoming inexpensive,accurate, and reliable and, with an average detection time onthe order of a few minutes, can significantly reduce the delaytime as well as bring the testing to doctors offices and patientshomes. As a result, the wide use of biomedical sensors maylead to more individualized health care services that will betailored for the needs of a patient and will match a specificgenotype. Indeed, one may envision using biomedical sensorsfor optimizing drug doses, monitoring the effectiveness oftreatments, and monitoring health conditions over personslife time. In addition to decreased time delays, miniaturiza-tion of biomedical sensors and integration with microfluidicdevices is yielding advanced analytical microsystems such asa BioChemLab-on-a-Chip. Integration of several sensors ona single substrate produces transducer arrays. A few recentexamples include electronic noses and tongues that are capableof performing multimeasurand detection in a few minutes.For a more complete listing of various biomedical, biological,and biosensing applications, the authors refer the reader tothe following references [138][140]. Overall, these excitingdevelopments forecast a biomedical sensor revolution thatcould dramatically change the way in which the medical,pharmaceutical, and environmental industries practice theirfields in the future.


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