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H* Journal of The Electrochemical Society, *** * H6-H11 2005

****-****/****/*** * /H6/6/$7.00 The Electrochemical Society, Inc.

Design and Testing of an Impedance-Based Sensor

for Monitoring Drug Delivery

Audrey M. Johnson,a Donald R. Sadoway,b,* Michael J. Cima,b

and Robert Langera,z

Department of Chemical Engineering and bDepartment of Materials Science and Engineering,

Massachusetts Institute of Technology, Cambridge, Massachusetts 02139, USA

A new impedance-based sensor to monitor drug delivery from an implantable microelectromechanical systems MEMS device

has been fabricated and tested. The sensor consists of two electrodes on opposing sides of a pyramidal drug reservoir. The

dissolution of the drug and advance of solution into the reservoir cause the impedance to change over time. A 100 times scale

model of the sensor was constructed to examine the effects of electrode geometry on solution resistance. An equivalent circuit was

formulated to interpret the impedance signal in terms of the resistance and double-layer capacitance of the solution in the reservoir.

The circuit was validated by impedance measurements on reservoirs lled with phosphate-buffered saline solutions of varying

concentrations. The sensor was then used to monitor the dissolution of the model drug mannitol from the drug delivery MEMS

device. The measured solution resistance and double-layer capacitance are related to the rate of transport of drug from the device,

making this sensor a potential instrument for noninvasive monitoring of drug transport from the implant in vivo. Experimental

results agree closely with the expected values of capacitance, resistance, and dissolution time calculated from physical parameters.

Manuscript submitted January 26, 2004; revised manuscript received June 11, 2004. Available electronically November 22, 2004.

imaging;17,18 however these techniques are limited in their resolu-

A sensor for characterizing the transport of a drug from an im-

tion, accuracy, and/or detection limits.19-21 Measurements of blood

plant in vivo, in real time and noninvasively, would have the poten-

tial to yield new information about the body s response to implants and urine drug levels also give some feedback on the operation of a

and the impact on drug or analyte transport. Understanding in vivo drug delivery device. However, many transport barriers exist be-

transport is critical for a number of applications including active tween the implant region and the point of measurement, and the

drug delivery devices and long-term implantable sensors. Most cur- complex pharmacokinetics of many drugs further complicates the

interpretation of these measurements.22-24

rent methods for measuring drug transport from implants are inva-

sive and destructive.1-4 We have designed and tested a novel sensor A drug delivery MEMS device has been developed previously in

that allows the monitoring of drug dissolution from an implantable our laboratory which allows the release of multiple drugs in a pul-

satile fashion.25,26 It consists of an array of microreservoirs etched

microelectromechanical systems MEMS device in real time with-

out disturbing the medium through which the drug is diffusing. into a silicon substrate, each capped by a thin gold membrane. Upon

Furthermore, the physical disposition of the sensor is such that application of a 1 V potential in the presence of chloride ion, the

without recalibration it can continue to give meaningful data even gold membrane oxidizes to form soluble gold chloride, thinning the

under changing conditions in the surrounding tissue, e.g., implant membrane until it fails. The drug within the reservoir then dissolves

encapsulation. into the surroundings as the aqueous solution advances into the

Implants are often walled off from the rest of the body by an reservoir.

avascular brous capsule of tissue that inhibits transport of sub- This paper presents a modi cation of the drug delivery MEMS

stances such as oxygen or nutrients between the body and the device in which two electrodes within each drug reservoir serve as a

implant.5 The formation of a brous capsule around implanted de- novel drug release sensor, as depicted in Fig. 1. Each reservoir has

vices has particularly important implications for drug delivery and the shape of a square pyramid due to the anisotropic etch used to

sensing applications. Current drug delivery implants are usually de- create them, and the electrodes cover two opposing sides of this

signed for a slow, steady release of a drug over time. In such a case, pyramid. Two-electrode impedance spectroscopy is used to measure

the transport pro le of drug is pseudo-steady state, so that while the the electrical characteristics of the reservoir. As the drug dissolves,

brous capsule affects drug release, the device is still able to the change in the electrical signature of the system allows real time

function.6,7 However, in the case of active devices, where quick monitoring of the rate of transport of drug away from the device.

pulsatile or more complex release pro les are desired, the isolation The utility of the sensor was demonstrated by in vitro measurements

of the implant causes a time lag due to the nite amount of time interpreted by means of an equivalent circuit.

required for diffusion of the drug through the capsule.8,9 This time

delay is also a key factor in the performance of in vivo sensors.10,11 Experimental

Even if sensor components are made resistant to the immune re- Chemicals. D-mannitol and potassium chloride Sigma, St.

sponse, the sensor must be continually recalibrated to account for Louis, MO were used as received. Phosphate buffered saline, PBS,

the changing conditions in the tissue surrounding it.12

consisting of 0.137 M NaCl, 0.001 M KH2 PO4, 0.01 M Na2 HPO4,

Most methods for determining the extent of implant encapsula-

and 0.0027 M KCl at pH 7.4 when diluted to 1 time concentration

tion and the drug dissolution pro le in the vicinity of an implant

Roche, Indianapolis, IN was purchased as a 10 times solution and

involve invasive methods that result in the destruction of the sample.

diluted to the appropriate concentration. Solutions were made with

Each data point requires a different implant and animal, contributing

deionized water passed through a MilliQ Millipore, Billerica, MA

to a high variability in measurements. Nondestructive methods have

system, resistivity above 17.8 M cm. Wires and bond pads were

signi cant limitations. Those involving direct observation of drug

insulated with EP42HT medical grade epoxy MasterBond, Hacken-

are restricted to unique systems such as the translucent rabbit ear13

sack, NJ, cured for 24 h at room temperature.

or rat dorsal skin clamped in a glass window for viewing.14 Phar-

macokinetic studies also involve monitoring of drug distributions in Microfabrication of devices. Drug delivery MEMS devices

vivo through such methods as microdialysis15,16 and nuclear were fabricated in the Microsystems Technology Laboratory at MIT.

Silicon wafers were 4 in., double-side polished, 300 m thick, 1-10

cm, p-doped 100 silicon WaferNet, San Jose, CA . A 100 nm

thermal oxide layer was grown on both surfaces and covered with a

* Electrochemical Society Active Member.

150 nm low pressure chemical vapor deposition LPCVD nitride

E-mail: abqq6a@r.postjobfree.com

Journal of The Electrochemical Society, 152 1 H6-H11 2005 H7

shadow mask was aligned over the device wafer Karl Suss MA4

Aligner, Suss MicroTec Inc., Waterbury Center, VT and 300 nm

gold over a 10 nm titanium adhesion layer was deposited via elec-

tron beam evaporation to form the impedance electrodes. A 650 nm

thick oxide layer was deposited by plasma deposition and patterned

over the front sides of the wafers to insulate the trigger electrodes

from the environment. Wafers were then diced into 5 5 mm de-

vices on an automatic dicing saw 2060 blade, DAD-2H/6T dicing

saw, Disco Hi-Tec America, Inc., Santa Clara, CA .

Scale model of device reservoirs. A 100 times scale model of a

square pyramidal drug delivery device reservoir was fabricated from

acrylic by the MIT machine shop. The pyramid was of base length

4.8 cm, height 3.0 cm, and base angle 54.4 . One side of the pyra-

mid was removable to provide access to the pyramid interior. Gold

foil electrodes 25 m thick, 99.9 %, Sigma were cut to cover two

opposing interior sides of the pyramid with tabs extending from the

top of the model for electrical connections.

PBS solution conductivity measurements. The conductivities of

PBS solutions were measured using a two-electrode conductivity

cell Industrial Instruments, Inc., Cedar Grove, NJ fully immersed

in the solution to be measured. The impedance spectrum was re-

corded using a Solartron 1255B frequency response analyzer with a

SI1287 electrochemical interface. The impedance was measured for

frequencies between 102 and 106 Hz, suf cient to reveal the critical

point or notch in the Cole-Cole plot. The critical point was indepen-

dent of the magnitude of the ac excitation voltage. The cell constant

was determined using standard solutions of 0.01 D,c 0.1 D, and 1.0

D potassium chloride. All solutions were kept at room temperature,

23.5 C, and the conductivities of the standard solutions at this tem-

perature were interpolated from values given by Wu and Koch. The

measured value of the cell constant was then used to determine the conductivity of 0.01,

0.1, 0.5, and 1.0 times PBS solutions.

Impedance of reservoir scale model. The two-electrode imped-

ance of the 100 times scale model of a device reservoir was mea-

sured for frequencies between 10 and 106 Hz. An initial aliquot of

1.0 times PBS was added to the model to make contact between the

Figure 1. The drug delivery MEMS device: a cutaway perspective view of foil electrodes and the impedance measurement was repeated. The

the prototype, with electrodes on the top surface for drug release and elec- impedance measurement and addition of PBS was repeated in incre-

trodes on the bottom surface and in the reservoirs for drug release monitor- ments of 1 to 2 mL solution until the model was entirely full. The

ing, dimensions 5 5 0.3 mm; b view of one device reservoir showing experiment was repeated using 0.5 times PBS and 0.1 times PBS

electrode con guration to scale; c idealized representation of drug release solutions.

from a reservoir.

Impedance of microdevice reservoirs. A drug delivery MEMS

device was packaged with open reservoirs for measurement of im-

pedance submersed in PBS solutions. The impedance electrodes

layer to act as an etch stop standard batch diffusion furnaces, were connected by traces to bond pads along the edges of the de-

Thermco 10K, La Porte, IN . The back sides of the wafers were vice. Insulated gold wires were epoxied MasterBond EP42HT

patterned with an array of 480 m squares using positive photoresist alongside the device and gold wire bonds, diameter 25 m, were

OCG825-20, Arch Chemicals, Norwalk, CT, which were plasma made between the bond pads and the wires. The wires, bonds, pads,

etched through the nitride and oxide layers Plasmaquest Series II and traces were covered with epoxy to insulate them. The two-

Reactor model 145, MKS Instruments, Andover, MA . Wafers were electrode impedance of each of four reservoirs was measured for

immersed in 5.17 M potassium hydroxide at 85 C until the silicon frequencies between 10 and 106 Hz. The device was then submersed

had been etched through the entire thickness of the wafers to form in PBS solutions of 1.0, 0.5, and 0.1 times concentration. The solu-

an array of pyramidal reservoirs capped by nitride/oxide mem- tions were observed to entirely ll the reservoirs, and the impedance

branes. The front sides of the wafers were then patterned with elec- measurement across each reservoir was repeated in each solution.

trodes for the triggering of drug release from reservoirs using image-

Impedance during drug release. The reservoirs of a drug deliv-

reversal resist AZ5214 E, Clariant, Somerville, NJ . Electron beam

ery MEMS device were completely lled with 35-45 g solid man-

evaporation model VES 2550, Temescal Semiconductor Products,

nitol per reservoir by a melt process. First, powdered mannitol was

Fair eld, CA was used to deposit 300 nm gold sandwiched between

spread over the reservoirs of a template device and heated, melting

two 10 nm titanium adhesion layers, and the image-reversal resist

it into the reservoirs. Excess mannitol was removed with a razor

pattern was lifted off in acetone. The remaining nitride and oxide

blade, and the pyramid-shaped mannitol pieces remaining were re-

were removed from the back sides of the wafers by plasma etching,

moved from the reservoirs. These presized mannitol pieces were

and a conformal 100 nm oxide layer was deposited by plasma depo-

then placed into the reservoirs of the drug delivery MEMS device,

sition at 80 C Plasmaquest Series II, model 145 . A shadow mask

wafer27 was prepared separately by deep reactive ion etching heated to melt them in place, and weighed AD-4 Autobalance,

DRIE etching a wafer with the pattern of the gold impedance

electrodes to be deposited inside the device reservoirs DRIE Mul- c

The demal D is a concentration unit used in connection with the electrical con-

tiplex ICP, Surface Technology Systems, Portsmouth, NH . The ductivity of aqueous solutions.28

H8 Journal of The Electrochemical Society, 152 1 H6-H11 2005

Table I. Conductivity of PBS solutions.

PBS concentration Calculated conductivity Measured conductivity

at 25 C S cm 1 at 23.5 C S cm 1

times

2 2

1.0 1.59 10 1.57 10

3 3

0.5 8.37 10 8.32 10

3 3

0.1 1.81 10 1.79 10

4 4

0.01 1.90 10 1.88 10

Perkin-Elmer, Boston, MA . The device reservoirs were then sealed

with a small piece of a glass coverslip and epoxy, and gold wires

were epoxied beside the bond pads. Wire bonds were made between

the impedance electrodes of each reservoir and the wires, then insu-

lated with epoxy. The two-electrode impedance of each reservoir

was measured at frequencies between 103 and 106 Hz. The device

was then submersed in 1.0 times PBS solution, and the impedance

measurement was repeated. After each reservoir was opened, the

reservoir impedance was measured every 3 min until the impedance

did not change with time. At this point it was assumed that the

dissolution of the mannitol was complete.

Results and Discussion

Figure 3. The dependence of the measured solution resistance in the

PBS solution conductivity. The measured values of the electri- square pyramidal model on height of the solution in the model and PBS

cal conductivity of solutions of 1.0, 0.5, 0.1, and 0.01 times PBS at concentration.

room temperature, 23.5 C, are given in Table I. For comparison, the

expected conductivity at 25 C was estimated by summing the molar

conductivities of the ionic species present in each PBS solution. Cole plot for such a system is a vertical line with a real intercept

Molar conductivities of ionic species were interpolated from re- equal to the solution resistance. In physical systems it is generally

ported values KCl, NaCl, except for 1.0 times PBS, which was observed that a constant phase element CPE ts the data better

extrapolated from the nearest reported value or extrapolated from than a simple capacitor. This is also true in our case, as the Cole-

Cole plot has a nite, positive slope. However, although it is gener-

the limiting molar conductivity (Na2 HPO4, KH2 PO4 ) using Kohl-

rausch s law.29,30 Standard KCl solutions gave a value of 1.00 ally assumed that some sort of dispersion of physical properties

0.01 cm 1 for the cell constant. gives rise to CPE behavior, the physics behind this are not well

understood,32 and we found it unnecessary to resort to the inclusion

Impedance of 100 times scale model. The scale model of the of a CPE in order to obtain a reasonably good t to the data.

pyramidal reservoir was constructed to evaluate the effects of the The resistance of a square pyramid can be calculated by starting

unusual electrode geometry on the impedance. The expected equiva- with the formula for conductance of a solid

lent circuit and impedance for the system, an aqueous solution with

no Faradaic processes occurring,31 is shown in Fig. 2. The Cole- Y A/ L

where A is the cross-sectional area, is the resistivity, and L is the

length. By taking a differential horizontal cross section, we can in-

tegrate the conductivities of each slice from x 0 to x H, where

H is the height of the pyramid. In this case, A and L are both

functions of x and of the base angle of the pyramid. For a square

pyramid, at any height, x, the width of the electrode at the edge and

the distance between the electrodes are the same function of x, L ( x ).

This means that the differential area is given by

dA L x dx

Substituting into the equation for Y gives the formula

dY L x dx/ L x

The L ( x ) cancel out, and the equation is easily integrated from zero

to H, yielding

Y H/, R /H

From this formula we see that for the special case of a square

pyramid, the resistance scales with the inverse of the height of the

pyramid. Therefore, as the pyramid model is lled with solution, the

resistance should scale with the inverse of the height of liquid in the

model. This matches the experimental ndings as shown in Fig. 3.

In an ideal case the slope of the plot should be equal to the solution

Figure 2. Equivalent circuit and Cole-Cole plot of experimental data and

resistivity, which is the inverse of the conductivity measurements in

best t to circuit for 100 times scale model of a square pyramidal drug

Table I. For our system the linear least squares slope for each solu-

reservoir containing 10 mL of 1.0 times PBS.

Journal of The Electrochemical Society, 152 1 H6-H11 2005 H9

Table II. Solution resistance and double-layer capacitance of mi-

croreservoirs lled with PBS.

Predicted Measured Measured

solution solution double layer

PBS concentration resistance resistance capacitance

times k k nF

1.0 2.13 1.79 0.12 3.54 0.79

0.5 4.01 3.68 0.44 2.50 0.62

0.1 18.6 16.0 4.7 1.9 1.6

This is likely due to fringing at the edges of the electric eld, which

is more signi cant in smaller systems. Taken together, these obser-

vations demonstrate the validity of the equivalent circuit of Fig. 4,

giving us a good understanding of the physical system.

Monitoring drug release by impedance measurements. A goal

of this work was to demonstrate that the measurement of impedance

would allow us to monitor the rate of drug release from a MEMS

device reservoir in real time. A typical example of the impedance

measurement during release of mannitol, a model drug, from a de-

vice reservoir is shown in Fig. 5a and b, as a succession of Cole-

Cole and Bode plots. The impedance before opening the reservoir is

Figure 4. Equivalent circuit and Cole-Cole plot of experimental data and similar to that observed for reservoirs full of air, which is expected

best t to circuit for a drug delivery MEMS device reservoir lled with 0.1 because mannitol is a semicrystalline solid with negligible conduc-

times PBS. Similar results were obtained for solutions of higher ionic

tivity, and there is no electrolyte present for conduction. After open-

strength and for different device reservoirs.

ing the reservoir, the impedance changes gradually from the single

semicircle characteristic of the RC parallel circuit to the overlapping

double semicircle characteristic of the circuit in Fig. 4. The solution

tion differs from the resistivity by a factor of 0.95 0.01, which resistance drops as the saline solution provides an ever larger low

gives us an effective cell constant for this particular experimental resistance path between the electrodes, and the double-layer capaci-

setup of 0.95/H, very close to the theoretical cell constant of 1/H tance increases as the area of the electrodes in contact with solution

derived above. The experimental data agree very well with the pre- increases. Successive data ts to this circuit give the values of so-

dicted equivalent circuit and expected values for solution resistance. lution resistance and double-layer capacitance vs. time shown in Fig.

In addition, the double-layer capacitance for an aqueous solution is 5c. This corresponds to the dissolution of the mannitol and the ad-

expected to be on the order of 10 F/cm2, and the best t capaci- vance of the PBS solvent front into the reservoir. The dissolution is

tance for our data agrees with this estimate. largely complete after 90 mins. Further experiments with radio-

labeled mannitol indicate that the gradual change in sensor output

Impedance of drug delivery MEMS reservoirs. The resistance

parallels the rate of release measured by scintillation counting of the

of a square pyramid calculated above was then compared to the

release medium.

resistance measured in the drug delivery device reservoir. For the

For comparison, a rst-order estimate of the expected time for

microsystem the equivalent circuit is not as simple, as shown in Fig.

dissolution can be made by considering the dissolution and diffusion

4. The parallel resistance and capacitance arise from a parallel cur-

of a solid in one dimension.33,34 A mass balance gives the following

rent path through the silicon substrate in which the reservoir is

equation for C ( x, t )

etched. This path is due to leakage current from the electrodes

through the silicon oxide layer beneath into the silicon, and takes the 2

C / x 2;

C/ t D C x,0 0

form of a parallel RC circuit, a so-called leaky capacitor. The mag-

nitude of the leak current is determined by the connection between C 0,t C sat ; C,t 0

the wire bonds and the bond pads. This varies from reservoir to

reservoir due to the inconsistency of the wire bonding process. For where C sat is the solubility of the compound in solution and D is the

any particular reservoir the resistance and capacitance of this path diffusivity. This can be solved for the ux at the interface as a

were found to be independent of reservoir contents, as expected. The function of time35

resistance remained constant within 3% during all experiments and

1/2

Nx D dC / dx C sat D / t

the capacitance was constant within 10%. x0 x0

The Cole-Cole plot of this circuit is two overlapping semicircles,

which corresponds well to the experimental data, as depicted in the A mass balance at the interface between solid and liquid can be

representative data t given in Fig. 4. The solution resistance de- written as

pends on the ionic strength of the solution and determines the posi-

A D dC / dx x 0 dt A C s C sat d z

tion of the in ection point between the semicircles, while the silicon

resistance depends on the device reservoir and determines the over-

where A is the area exposed to solution, C s is the density of the solid

all width of the curve. Some depression of the data from the ideal t

was observed, but again, a reasonably good t was obtained without compound, and dz is the differential distance traveled by the inter-

the substitution of a CPE for the capacitive elements. face in the time dt . Note that the equation for C ( x, t ) is only valid if

The average solution resistance and double-layer capacitance of we make the assumption that the motion of the interface is slow

the device reservoirs are given in Table II. The double-layer capaci- compared to the formation of the concentration pro le pseu-

tance is of the correct order of magnitude, a factor of 104 smaller dosteady state . If we substitute the expression for the ux at the

than that observed for the 100 times scale model. The measured interface into the interfacial mass balance, we obtain an expression

solution resistance is lower than that calculated using the formula for the distance traveled by the interface the depth of penetration of

solvent into the reservoir as a function of time

for the resistance of a square pyramid by a factor of 0.86 0.02.

H10 Journal of The Electrochemical Society, 152 1 H6-H11 2005

Figure 5. Impedance of a MEMS de-

vice reservoir during release of manni-

tol into 1.0 times PBS. a Series of

Cole-Cole plots over time, b series

of Bode plots over time, and c series

of best ts of solution resistance and

double-layer capacitance over time.

allow repeated measurements of drug release from the same implant

1/2

zt 2 C sat / C s C sat Dt/

at multiple time points, without the need for explantation or disrup-

tion of the tissue environment.

Conclusion

Substituting appropriate values for the physical parameters ( C sat

0.18 g/cm3, C s 1.5 g/cm3, D 10 5 cm2/s, z 300 m, The drug delivery MEMS device with an impedance-based sen-

we obtain an estimated time for dissolution of 80 min, which sor successfully monitored the release of a drug from a device res-

agrees quite closely with the experimentally observed dissolution ervoir. The system was represented by a simple equivalent circuit

time. whose elements correspond to the physical characteristics of the

The impedance-based sensor was used to monitor dissolution of system. The solution resistance and double-layer capacitance were

mannitol in vitro and the relationship between sensor output and the measured as functions of time during release of the drug. These two

physical system was characterized by an appropriate equivalent cir- parameters are functions of the degree of penetration of solution into

cuit. Future experiments will evaluate the relationship between the the reservoir during drug release, which can be related to the rate of

sensor output and the rate of transport of the drug through various transport of the drug from the device. The sensor has the potential to

transport barriers. In addition, it would be useful to compare sensor be used as a novel method for noninvasive monitoring of drug trans-

output to traditional methods of evaluating drug release, such as port from the implant in vivo.

radio-label or uorophore detection. Finally, a critical test will be to

Acknowledgments

use the sensor to monitor drug release during long-term implantation

in vivo, evaluating the impact of brous capsule development on We thank Dr. D. DeLongchamp and Professor M. Zahn for dis-

drug transport. The nondestructive nature of the measurement will cussions of equivalent circuit analysis, and Dr. R. Shawgo and Dr. L.

Journal of The Electrochemical Society, 152 1 H6-H11 2005 H11

Arana for help with packaging and microfabrication aspects of the 16. M. Muller, Adv. Drug Delivery Rev., 45, 255 2000 .

17. K. A. Salem, A. Szymanski-Exner, R. S. Lazebnik, M. S. Breen, J. Gao, and D. L.

project. This work was supported by U.S. NIH grant no. 1 R24

Wilson, IEEE Trans. Med. Imaging, 21, 1310 2002 .

AI47739-01 and graduate fellowships from the U.S. NSF and

18. A. Szymanski-Exner, N. T. Stowe, K. Salem, R. Lazebnik, J. R. Haaga, D. L.

Merck. Wilson, and J. Gao, J. Pharm. Sci., 92, 289 2003 .

19. W. Wolf, Adv. Drug Delivery Rev., 41, 1 2000 .

Massachusetts Institute of Technology assisted in meeting the publication

20. R. E. Port and W. Wolf, Invest. New Drugs, 21, 157 2003 .

costs of this article.

21. A. J. Fischman, N. M. Alpert, and R. H. Rubin, Clin. Pharmacokinet., 41, 581

2002 .

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