Abstract— Tightly controlling blood glucose in selected

patients has been shown to significantly decrease complications

such as infection, neurologic injury, and respiratory failure.

Intravenous microdialysis has been used to continuously

monitor glucose levels yet variable recovery may occur due to

probe fouling, osmosis, and ultrafiltration. A portable

microdialysis system is presented that: (1) uses retrodialysis

with fluorescein to monitor probe fouling, (2) measures

downstream flow rate to gauge fluid losses from

ultrafiltration/osmosis, and (3) employs a wireless

microcontroller to sequence the above events and

record/transmit the data. In vitro glucose microdialysis is

performed with a CMA Microdialysis IView catheter and

Medtronic CGMS® iPro™ glucose sensor. Providing

self-diagnostics can eventually lead to a medical embedded

system with calibration, safety alarms, and wireless

communication of blood glucose readings to a drug delivery

system for closed loop control.

I. INTRODUCTION

ICRODIALYSIS is a versatile technique for sampling

molecules (analytes) from tissues or blood in living

animals and man. Microdialysis is performed by circulating a

small volume of buffer through a porous polymeric catheter

(probe) that is about 500 microns in diameter and 1 – 4 cm in

length. Pores in the probe allow the diffusion of molecules

from blood or the tissue of interest into the circulating buffer.

The buffer exiting the probe, also known as the dialysate, is

then collected for later analysis. Microdialysis has been used

to continuously measure glucose levels in the subcutaneous

tissue of diabetics [1]-[3].

Recently, commercially manufactured microdialysis

probes have become available for continuous blood glucose

monitoring (CMA Microdialysis, Chelmsford, MA, U.S.A.).

Researchers have demonstrated that tight blood glucose

control in critically ill patients has multiple benefits [4]. Such

control is achieved by continuous intravenous insulin

infusion. It is not without certain risks; subgroups of patients

Manuscript received February 15, 2009. This work was supported in part by

the the Commonwealth of Pennsylvania under A Pennsylvania

Infrastructure And Technology Alliance (PITA) Grant.

A. J. Rosenbloom is with the Institute For Complex Engineered Systems,

Carnegie Mellon University, Pittsburgh, PA. 15213 USA; Phone:

412-268-8638; fax: 412-268-6571; e-mail: alan2@andrew.cmu.edu.

H.R. Gandhi is with the Electrical and Computer Engineering Dept.,

Carnegie Mellon University, Pittsburgh, PA. 15213 USA; Phone:

412-268-5212; e-mail: hgandhi@andrew.cmu.edu.

G.L. Subrebost is with the Institute For Complex Engineered Systems,

Carnegie Mellon University, Pittsburgh, PA. 15213, Carnegie Mellon

University, Pittsburgh, PA 15213 USA; Phone: 412-268-5212; e-mail:

georgel@andrew.cmu.edu.

are actually harmed by tight control, perhaps by the ill effects

of hypoglycemia [5]. Blood glucose must be followed

closely in order to keep levels within 80 to 140 mg/dL.

Measurements are normally taken every 1 – 2 hours to adjust

the insulin dose properly. Continuous blood glucose

monitoring via microdialysis would eliminate repeated

testing and allow adjustment of glucose level more smoothly,

quickly and with increased safety. However, there are factors

that potentially create variability with estimates of blood

glucose obtained by microdialysis. Fouling of the probe, i.e.

clogging of its pores by blood proteins or clotting, can

decrease the recovery of glucose. Furthermore, hydrostatic

or osmotic pressure differences between the microdialysis

buffer and the blood or tissue being sampled will cause a net

flow of fluid across the probe membrane (i.e. ultrafiltration

or osmosis of water). Ultrafiltration into the probe will

augment analyte recovery while ultrafiltration out of the

probe will reduce recovery. The importance of ultrafiltration

increases with membranes having larger pore sizes [6].

Microdialysis modeling has been used to provide more

insight to the complex trans-membrane dynamics [7].

The variability in analyte recovery caused by clogging,

osmosis, and ultrafiltration can become important when

potentially harmful therapy will be based on glucose

readings. Furthermore, intravenous microdialysis systems

can fail for many reasons such as the probe pulling out of the

vein, broken connections or fluidic lines, clotting in the vein,

etc. In order for intravenous microdialysis to move toward

being a robust clinically useful modality, there must be

alarms to warn of system failure and self-checks to confirm

adequate function. These capabilities would ideally be

integral to the microdialysis setup. In this paper we

demonstrate a microdialysis system that performs continuous

glucose measurements using a commercially available

glucose electrochemical sensor. We explore methods to

make the system self-monitoring.

The two monitoring methods are: (1) detection of flow rate

after the microdialysis probe – this will confirm pump

function and system integrity, and detect excess fluid being

taken from or added to the system by ultrafiltration or

osmosis, (2) modified retrodialysis to assess probe patency

and membrane diffusion capacity. Retrodialysis is normally

performed by adding a marker molecule very similar to that

being sampled to the dialysis buffer in order to calibrate the

recovery of analyte [8]. The diffusion rate of the marker out

of the probe is semi-quantitative for determining the

diffusion rate of the similar analyte into the probe. This and

other methods of calibration [9] are particularly useful in

tissue microdialysis where the kinetics of transport of

analytes into the microdialysis probe are very complex.

2009 IEEE/ICME International Conference on Complex Medical Engineering

Glucose Microdialysis with Continuous On-Board Probe

Performance Monitoring

Alan John Rosenbloom, Heer Robin Gandhi, George Lopez Subrebost

M

978-1-4244-3316-2/09/$25.00 ©2009 IEEE

These methods are less applicable with bloodstream

microdialysis of small molecules such as glucose. Due to a

continuous blood flow that mimics sampling from a stirred

solution, the diffusion rate of small molecules is very rapid,

allowing good equilibration with the microdialysis buffer.

Thus, the principal barrier to analyte recovery is fouling

(clogging) or clot formation on the probe membrane. We

confirm probe patency by continuously quantitating

fluorescein, a fluorescent molecule. Numerous researchers

have demonstrated miniature fluorescence detection systems

with various dyes [10-12]. Fluorescein is approved for

intravenous use in humans and has the potential for use in

clinical settings. We explore continuous flow rate

determination after the probe to quantitate ultrafiltration,

confirm fluidic transport within the system and to detect

system failures such as kinked or broken lines or

connections. By using ‘timing marks’, created by

photobleaching fluorescein present in the microdialysis

buffer and tracking this bleached ‘plug’ further downstream,

time-of-flight calculations are used to determine volumetric

flow rate [13]. Alternate methods such as ultraminiature

thermodilution sensors can also be used.

II. METHODS AND MATERIALS

A. Optofluidic Setup

A CMA 107 syringe pump (CMA Microdialysis,

Chelmsford, MA, U.S.A.) is used to provide various volume

flow rates from 0.1 to 10 µL (see Figure 1). In vitro glucose

sampling was performed with an IView CMA 64

microdialysis catheter (also manufactured by CMA

Microdialysis). This catheter is intended for continuous

intravenous monitoring of hospitalized patients during

surgery, intensive care or in the general wards [14]. With a

membrane cut-off of approximately 20,000 Daltons, the

catheter can also be used to measure free fractions of drugs in

blood during pharmacokinetic and pharmacodynamic

studies. The porous membrane material is made of

polyarylethersulphone (PAES) with an outer diameter of 0.6

mm and length of 10 mm. The outlet tubing of the catheter

(inner diameter 120 µm), normally connected to microvials

for fraction collection of the dialysate, was instead

lengthened to accommodate the monitoring schemes

mentioned above by attaching Tygon tubing (inner diameter

508 µm).

Downstream from the microdialysis catheter, a

photobleaching and fluorescence detection region has been

assembled to perform fouling and ultrafiltration

measurements. Photobleaching of the fluorescein in the

dialysate is achieved by locating the dialysate tubing within a

beam of focused light from a light emitting diode (LED),

model Luxeon V, LXHL-LB5C (Philips Lumileds, San Jose,

CA, U.S.A.). This LED has a Lambertian radiation pattern,

Fig. 1. Continuous glucose microdialysis setup that uses a CMA Microdialysis IView CMA 64 microdialysis catheter and Medtronic CGMS®

iPro™ glucose sensor. An embedded systems microcontroller is used to control the photobleaching as well as monitor the fluorescence of the

dialysate.

Fig. 2. (Top) Medtronic CGMS® iPro™ data recorder and sensor probe

used to continuously record dialysate glucose values. (Bottom)

Amperometric readings from the Medtronic CGMS® iPro™ glucose sensor

placed inside a tube with inner diameter of 794 µm and connected to a 10

mL syringe. Large dips between reading plateaus are due to manual

flushings from the syringe.

outlet luminous flux of 48 lumens, and has a dominant

wavelength of 470 nm, which is necessary for exciting

fluorescein. A narrow beam lens (Model FLP, Fraen Corp.,

Reading, MA, U.S.A.) with an aspheric profile, and total

beam divergence of 12 degrees, is mounted over the LED to

provide a collimated light source. Another narrow beams

lens is placed in front of the collimated light (forming an oval

enclosure) to focus the beam to the dialysate tubing for

photobleaching. Due to significant heating, the LED is

mounted onto an aluminum heat sink to provide adequate

temperature regulation.

A miniature fluorometer setup (8 cm x 4 cm x 2.5 cm) was

made from a machined plastic block (acrylonitrile butadiene

styrene) to measure fluorescence of the dialysate as well as

the photobleaching front. A collimated light source is

provided by another Luxeon V LED and FLP series narrow

beam lens. To limit the output light from corrupting the

fluorescence detection, the collimated light is filtered by an

exciter (ET470/40x, Chroma Tech. Corp., Rockingham, VT,

U.S.A). An aspheric lens (Model 352330-A, ThorLabs Inc.,

Newton, NJ, U.S.A.) with a diameter of 6.35 mm, focal

length of 3.1 mm, and numerical aperture of 0.68 is used to

focus the light to a 0.5 mm spot on the dialysate tubing.

Fluorescence emission is detected orthogonally from the

light source path. A second aspheric lens with the same

specifications is used to magnify the emitted light output.

This output is filtered with an emitter (ET525/50m, Chroma

Tech. Corp.) and collected by a silicon photodiode (BPW21,

Vishay Semiconductor, Heilbronn, Germany). The output

signal from the photodiode is conditioned and amplified by

an embedded systems microcontroller that will be further

described.

B. Glucose Sensing and Reagents

100 µM fluorescein in 1X calcium magnesium free –

phosphate buffered saline (CMF-PBS) was used as the

source buffer for the syringe pump. Fluorescein (MW 332

g/mol) was obtained from Sigma Aldrich (St. Louis, MO,

U.S.A.). To test the response of the photodiode, eight

different fluorescein concentrations were prepared from 5 to

85 µM.

The microdialysis probe was placed in a 15 mL test tube

that contained glucose at a concentration of 180 mg/dL in a

1X CMF-PBS buffer. Anhydrous dextrose (glucose, MW

180 g/mol) was obtained from Fisher Scientific (Pittsburgh,

PA, U.S.A.).

An electrochemical glucose sensor, CGMS® iPro™

(Medtronic Inc., Northridge, CA) was used to continuously

record dialysate glucose values (see Figure 2a). The outer

diameter and total length of the sensor tip is approximately

650 µm and 2 cm, respectively. The sensor is typically used

by diabetic patients to monitor subcutaneous glucose levels

and provides a reading every 5 minutes. A calibration was

performed to determine the linearity of the glucose readings

(see Figure 2b). The glucose sensor is placed downstream

from the microdialysis probe inside a larger diameter Tygon

tubing (inner diameter of 794 µm) in order to accommodate

the larger size of the sensor tip. Glucose readings are stored

on-board using a recorder that attaches to the backside of the

probe and can be subsequently downloaded to a computer

using a wireless connection.

C. Electrical Setup

An embedded systems microcontroller (MSP430, Texas

Instruments, Dallas, TX, U.S.A.) is used to control the

sequence of photobleaching and fluorometer measurements

(see Figure 3). A custom printed circuit board, attached to the

microcontroller, is used to efficiently amplify the photodiode

signal and determine the fluorescence level in the dialysate.

Its secondary tasks are to sustain itself in terms of power

usage as well as retaining or transmitting the data that it

collects from the photodiode. To perform these tasks the

electrical setup is governed by a central microcontroller that

controls the functioning of the complete system. The system

is divided into three functioning modules.

1) Analog Amplification and Filtering

Fluorescence from the dye emits green light at a 521 nm

wavelength when blue light (480 nm wavelength) is

projected on to the dialysate tubing. The photodiode is

subjected to this emission radiation as well as other external

noise sources (60 Hz noise from light sources such as

fluorescent light tubes). To suppress this noise, the source

LED light is modulated to provide an identifier for

subsequent filtering and amplification. The fundamental

signal frequency chosen for this operation is 965 Hz and an

active band-pass analog filter (2nd

order, signal gain 40 dB) is

implemented using discrete off-the-shelf components. The

signal gain is programmable using a digital potentiometer so

that the central microcontroller can calibrate the gain of the

system based on the input signal level. This prevents the

signal from saturating if the optofluidic setup is altered. The

amplified and filtered analog signal is sent over to the

on-chip analog to digital converter (ADC) on the central

microcontroller.

2) Digital Calibration and Signal Level Detection

To measure the fluorescence level digitally the LED

generates a 965 Hz signal that is focused on the dialysate

tubing for approximately 500 ms. Within this time frame, the

ADC is programmed to sample the amplified signal for 2,560

samples. The average DC level and AC RMS level of the

signal are measured using digital computations on the

microcontroller. The analog gain can be calibrated using this

information and all the data henceforth is measured relative

to the calibrated signal. The AC RMS signal level is what

denotes the fluorescence level.

3) Storing and Transmitting Data

The electrical system is powered by a 3.7 V, 400 mA

lithium ion battery. Two dedicated ADCs are used to

measure the battery voltage and input charging voltage. The

microcontroller charges the battery from the external

charging voltage when the battery level is low using a voltage

supply of 4.5-5.5 V. Based on the power consumption of the

electrical setup, it should be able to sustain itself on battery

power for at least 5 days. The AC RMS values of the signal

are stored in the data memory of the microcontroller. A

CC2500 2.4GHz RF chip, embedded along with the

microcontroller, transmits the data stored to a receiver node

where it can be viewed and archived.

III. EXPERIMENTAL RESULTS

Serial dilutions of fluorescein (from 5 to 100 µM) were

introduced into the setup in order to test whether the

photodiode was accurately measuring fluorescence levels as

shown in Figure 4. Calibration of the photodiode output is

performed by filling the dialysate tubing with CMF-PBS

buffer solely and setting this to 0 µM fluorescein. The

photodiode output recorded at 100 µM fluorescein was then

used to determine a conversion factor between voltage and

concentration. Each fluorescein concentration was

introduced for 40-60 seconds and the data was recorded by

the microcontroller every second during an LED pulse. The

data shows very good linearity as the concentration was

increased. While the fluorescein was stagnant within the

setup, a noticeable decline in fluorescence was observed.

Although the power supply on the LED was low (60 mA),

some inadvertent photobleaching occurred even though the

LED is only powered for 500 ms.

Fig. 4. Plot showing photodiode response with varying fluorescein

concentration. Calibration was done by using CMF-PBS buffer to set the 0

mM floor, while the 100 µM fluorescein photodiode output was used to

arrive at a voltage-to-concentration conversion factor.

In order to calculate volumetric flow rate in the dialysate

tubing, a photobleaching section was placed approximately

35 mm upstream from the fluorometer measurement. The

syringe pump was filled with 100 µM fluorescein and set to 2

µL/min. A tubing length of approximately 15 mm was

photobleached with a high LED intensity (500 mA) for 1

minute. Based on the inner diameter of the dialysate tubing

(508 µm), it should take 3.5 minutes for the photobleached

Fig. 3. Embedded systems controller includes a Texas Instruments

MSP430 microcontroller and a custom printed circuit board for signal

amplification and filtering.

plug to reach the fluorometer. As shown in Figure 5, this plug

takes about 3.7 minutes for the fluorescence intensity to

begin a steep decline caused by the photobleaching. The

difference in timing is probably due to inaccurate estimation

in length between the upstream bleaching and downstream

fluorescence measurement. A second photobleaching event,

as shown in Figure 5, was performed with a lower LED

intensity (275 mA) for 1 minute as well.

Fig. 5. Plot showing photobleached ‘plug’ passing the fluorometer

measurement region. A 12 mm section of outlet tubing was photobleached

upstream from the fluorometer. A high (500 mA) and low (275 mA) LED

power setting created the left and right dips, respectively.

Lastly, the Medtronic glucose sensor is combined with the

modified retrodialysis method in order to evaluate the

accuracy of using fluorescein as a metric for glucose

trans-membrane diffusion. The microdialysis catheter was

placed into a 15 mL test tube that contained a 180 mg/dL

glucose solution. Unfortunately, a mixing stirrer was not

placed inside the test tube to provide adequate refreshing of

the glucose solution. Before taking glucose readings, it is

necessary to allow for a 2 hour initialization time after the

sensor is wetted. A 100 µM fluorescein solution was used as

the microdialysis buffer while the volumetric flow rate on the

syringe pump was varied to 4 different intervals (2, 5, and 10

µL/min). A low LED intensity was used for the fluorometer

measurement (60 mA) and data was recorded every second.

The photobleaching setup was not used in this experiment.

Figure 6 shows two curves representing the relative

recovery of glucose and relative loss of fluorescein as the

flow rate was varied. A small flow rate causes the fluorescein

to leach out of the probe, while also providing adequate time

for the glucose to diffuse into the probe, yet this requires a

longer waiting period to perform the glucose sensing and

monitoring methods due to the length of the dialysate tubing.

High flow rate causes the fluorescein to bypass the diffusion

process out of the probe almost completely. A 10 µL/min.

flow rate shows a relative recovery of 80% fluorescein and

30% glucose, while a 2 µL/min. flow rate shows a relative

recovery of 35% fluorescein and 60% glucose. Since the

fluorescein molecule is approximately twice the size of

glucose (332 g/mol vs. 180 g/mol), it should provide an early

warning if fouling begins to occur on the probe membrane.

Further in vitro recovery tests will be performed using whole

blood before testing moves to a clinical trial with intensive

care unit patients. In addition, ultrafiltration will be induced

by altering the height of the dialysate tubing end, which is

exposed to atmospheric pressure. This should demonstrate

the adequacy of continuously measuring the fluorescein as a

gauge for probe fouling as well as measuring volumetric flow

rate to determine ultrafiltration values.

Fig. 6. In vitro glucose relative recovery and fluorescein relative loss at

varying volumetric flow rate. The microdialysis probe was placed in a test

tube containing 180 mg/dL glucose while the microdialysis buffer

contained 100 µM fluorescein.

IV. FUTURE APPLICATIONS

An embedded system microcontroller can be used to

perform the previously mentioned monitoring functions as

well as control the microdialysis flow rate. It is well known

that the slower the flow rate of microdialysis buffer, the more

complete the equilibration of analyte concentration between

sample fluid and buffer in the probe. However, very slow

flow rates impose a large delay in transporting the fluid to the

analyzer section of the system. Also, in-line analyzers require

a minimum sample volume. With sophisticated control of

microdialysis flow rate, one can achieve a highly efficient

system that collects sample at a slow rate for a set time. A

higher flow rate can be used to transport the sample to the

in-line analyzer. Finally, a slow rate can again be used to

allow time for the sample volume to reach steady state with

the analyzer components. We plan to integrate a servo

actuator to control of the syringe pump flow rate. The use of

an embedded systems microcontroller along with

microdialysis has other benefits as well. In the future, the

microcontroller can contain alarms, provide continuous

wireless communication of glucose levels, and communicate

with drug delivery systems to achieve closed loop control of

blood glucose. This technology is applicable to other drug

therapies as well as insulin control of glucose level.

ACKNOWLEDGMENT

The authors would like to thank CMA Microdialysis for

donating the IView CMA 64 microdialysis catheter.

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