An Electrochemical Micro-System for

An Electrochemical Micro-System for
Electrochimica Acta 163 (2015) 260–270
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An Electrochemical Micro-System for Clozapine Antipsychotic
Treatment Monitoring
Hadar Ben-Yoav a,b,1, * , Sheryl E. Chocron a,c, Thomas E. Winkler a,c , Eunkyoung Kim d ,
Deanna L. Kelly e, Gregory F. Payne c,d , Reza Ghodssi a,b,c, **
MEMS Sensors and Actuators Laboratory (MSAL), University of Maryland, College Park, MD 20742, United States
Department of Electrical and Computer Engineering, Institute for Systems Research, University of Maryland, College Park, MD 20742, United States
Fischell Department of Bioengineering, University of Maryland, College Park, MD 20742, United States
Institute for Bioscience and Biotechnology Research, University of Maryland, College Park, MD 20742, United States
Maryland Psychiatric Research Center, University of Maryland School of Medicine, Baltimore, MD 21228, United States
Article history:
Received 5 December 2014
Received in revised form 11 February 2015
Accepted 13 February 2015
Available online 17 February 2015
Clozapine is the most effective antipsychotic medication for schizophrenia, but it is underutilized
because of the inability to effectively monitor its treatment efficacy and side effects. In this work, we
demonstrate the first analytical micro-system for real-time monitoring of clozapine serum levels. An
electrochemical lab-on-a-chip is developed and integrated with a catechol-chitosan redox cycling
system. The microfabricated device incorporates 4 electrochemical reaction chambers with the
capability of analyzing microliter-volume samples. Integration of the catechol-chitosan film amplifies the
clozapine oxidative signal and improves the signal-to-noise ratio, which addresses sensitivity and
selectivity challenges. Optimization of the redox cycling system fabrication parameters and analysis of
various electrochemical techniques and data processing approaches is implemented to maximize
clozapine detection performance. The device is tested with buffer samples containing clozapine and
demonstrates a sensitivity of 54 mC mL cm2 mg1 and a limit-of-detection of 0.8 mg mL1, a sensing
performance similar to a counterpart macro-scale benchtop system. Importantly, the feasibility to
differentiate between 0.33 mg mL1 and 3.27 mg mL1 clozapine concentrations in human serum without
any preceding dilution or filtering procedures is demonstrated, a significant step towards utilizing pointof-care testing micro-systems for schizophrenia treatment management. With these micro-systems, we
envision more effective and safe treatment that will enable fewer visits to the clinicians, decrease costs
and patient burden.
ã 2015 Elsevier Ltd. All rights reserved.
Therapeutic drug monitoring (TDM)
Antipsychotic clozapine
Redox cycling system
1. Introduction
Schizophrenia is one of the most challenging and complex
neuropsychiatric disorders. It is a chronic illness that affects 1% of
the worldwide population [1,2]. While a lifelong treatment with
antipsychotics is recommended, approximately half of the
schizophrenia population is not receiving appropriate care [3].
* Corresponding author at: Institute for Systems Research, Department of
Electrical and Computer Engineering, University of Maryland, College Park, MD
20742, United States. Tel.: +1 301 405-2168; fax: +1 301 314 9920.
** Corresponding author at: Institute for Systems Research, Department of
Electrical and Computer Engineering, University of Maryland, College Park, MD
20742, United States. Tel.: +1 301 405-8158; fax: +1 301 314 9920.
E-mail addresses: [email protected] (H. Ben-Yoav), [email protected]
(R. Ghodssi).
ISE member.
0013-4686/ ã 2015 Elsevier Ltd. All rights reserved.
Clozapine (CLZ) is the most effective antipsychotic treatment for
chronic and treatment refractory patients with schizophrenia
[3–8]. It is one of the few antipsychotics whose efficacy has been
predicted by blood measurement [9–14]. An optimal therapeutic
range (0.350–1 mg mL1) has been identified for high efficacy and
low toxicity [13]. Current pharmacologic guidelines recommended
that CLZ blood levels be monitored for optimal therapeutic
response [3]. In spite of its superior effectiveness and the
recommended guidelines, CLZ is under-prescribed and underutilized in the United States [15–17]. In fact, CLZ is suggested to be the
most under-used evidence-based treatment in psychiatry [18–21].
This underutilization is proposed to be related to the burden to the
patients of repeated blood draws (weekly for the first six months)
to monitor for agranulocytosis (severely lowered white blood cell
count), a rare but dangerous side effect [22–24]. Furthermore,
nonadherence to medications is a widespread problem, and is one
of the more frequent reasons for relapse, re-hospitalization and
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
higher treatment costs in people with schizophrenia [25].
However, real-time monitoring of CLZ blood levels can be used
to optimize treatment efficacy, reduce the risk of toxicity, and
allow physicians to know if patients are adherent to their
treatment or need interventions to support adherence.
Electrochemical lab-on-a-chip (LOC) sensing micro-systems
provide numerous advantages in clinical diagnostics, environmental monitoring, and biomedical research fields [26–30]. These
micro-systems have been rapidly adopted by a number of clinical
diagnostics applications bringing benchtop methods to the pointof-care (POC) [31]. LOC devices have rarely been utilized in mental
health applications [32–39]. The few electrochemical-based
examples including neurotransmitter measurement [40,41], stress
analysis [42], oxidative stress assessment [43], and antipsychotic
drug monitoring [44]. In recent years, the electrochemical activity
of CLZ [45,46] and its detection [47–53] have been investigated
with macro-scale benchtop systems, a minority of which have
analyzed CLZ in blood samples [45,53–57]. When translating
analytical macro-scale technology into the micro-scale regime, a
number of challenges arise. Sample preparation and mixing with
other fluids, physical and chemical effects (e.g., capillary forces,
surface roughness, chemical interactions of construction materials
on reaction processes), and low signal-to-noise ratio are among the
most important problems to overcome in analytical micro-systems
[58]. Furthermore, the sensitivity and the specificity of these
micro-systems when testing complex biological fluids (e.g.,
cerebrospinal fluid, serum, saliva, and urine) are deteriorated
due to interference of other electro-active analytes and nonspecific adsorption to the surface of the electrode [59]. These
challenges may affect the performance of the device and limit its
ability to sense low drug concentrations in biological fluids.
Utilizing LOC technology for therapeutic drug monitoring (TDM) of
CLZ in schizophrenia will provide frequent, real-time blood levels
to maximize efficacy and minimize toxicity [22]. Such information
will enable clinicians to manage CLZ treatment and to close the
healthcare loop between clinicians and patients. No studies to date
have developed miniaturized micro-systems designed for TDM of
CLZ at the POC.
In this work, an electrochemical LOC for real-time monitoring of
CLZ blood levels is presented. The device is designed for minimal
preparation steps and utilizes a redox cycling system for CLZ signal
amplification (Scheme 1). Lithography and thin film metal
deposition technology is used to fabricate the device. The
fabricated device is modified with a thin film consisting of the
redox cycling system. This modification is based on a 2-step
biofabrication process that entirely relies on electrical signals
(i.e., chitosan electrodeposition and catechol electro-modification)
(Scheme 1A). In these processes, phenolic molecules (e.g., catechol
[60], melanin [61], and lignin derivatives [62]) are being used as
the electronic materials for bio-electronic devices. Specifically, the
catechol-chitosan redox cycling system has been demonstrated to
detect CLZ [53,63–66] with capabilities for miniaturization and
system integration [38,67–69]. It has also shown to be a stable film
with a small signal loss due to slow overall biomaterial degradation
[70]. With the catechol-chitosan redox cycling system, CLZ is
oxidized on the surface of the electrode and continuously
regenerated back to its reductive state (Scheme 1B; top left
figure). This redox cycling mechanism significantly increases the
generated CLZ oxidative current, amplifying CLZ electrochemical
signal. Recharging of the redox cycling system is accomplished by
the introduction of a ruthenium compound, which intermittently
reduces the catechol into its original state (Scheme 1B; top right
Here, we present the electrochemical validation and film
optimization of the device, as well as the functionality and the limit
of detection (LOD) of CLZ sensing. Additionally, we demonstrate
the feasibility of differentiating between 0.33 and 3.27 mg mL1
CLZ concentrations in undiluted and unfiltered human serum, a
first step towards reducing the burden and allowing more frequent
CLZ blood monitoring for the patient and clinicians.
2. Experimental
2.1. Test solutions and instrumentation
All chemicals were purchased from Sigma–Aldrich. Phosphate
buffered saline (PBS; pH 7.4) was prepared by dissolving a
concentrated tablet with deionized (DI) water resulting in a
solution of 0.01 mol dm3 phosphate buffer, 0.0027 mol dm3
potassium chloride, and 0.137 mol dm3 sodium chloride. All
electrochemical tests were performed with a CHI660D single
channel potentiostat and an 8-channel multiplexer (CH
Scheme 1. Redox cycling system integrated in an LOC for CLZ detection. (A) Device/Film Fabrication. Catechol-chitosan redox cycling modified LOC fabrication methods
(fabrication of the electrochemical chip, electrodeposition of chitosan on the working electrode, and modification of the chitosan film with catechol moieties).
(B) Electrochemical Detection. 3-electrode electrochemical micro-chamber in the LOC modified with the catechol-chitosan redox cycling amplification film for CLZ sensing
mechanism (top left) followed by recharging of the film with ruthenium (top right).
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
Instruments, Inc.). The potential values presented in this manuscript are all in reference to a Ag/AgCl half cell potential.
Electrochemical validation of the device was performed by testing
a mixture of 10 mmol dm3 PBS with 100 mmol dm3 sodium
chloride along with 5 mmol dm3 ferricyanide ([Fe(CN)6]3) and
5 mmol dm3 ferrocyanide ([Fe(CN)6]4) yielding a solution with
the reversible redox couple ions. A 28 mL volume was used with all
solutions for testing with the device. The chitosan electrodeposition solution contained 1% dissolved chitosan in diluted HCl at pH
5–6. The catechol modification solution contained 5 mM of
pyrocatechol dissolved in PBS solution. Additionally, 25 mmol
dm3 of hexaammineruthenium(III) chloride (Ru3+; Ru) was
dissolved in all tested solutions that indicate Ru. A film
initialization step was carried out to fully reduce the chitosanbound catechol residues and to consume all unbound catechol
with a mixture of Ru and 25 mmol dm3 of 1,1’-ferrocenedimethanol (Fc) dissolved in PBS. Commercial human serum (from
human male AB plasma, USA origin, sterile-filtered) was divided
into 1 mL aliquots and stored in 20 C conditions. Prior to the
experiment, the serum was thawed by horizontally lying down the
tube on an ice bucket at room temperature.
2.2. Lab-on-a-chip design and fabrication
The design of the electrochemical LOC used in this work
comprises four independent reaction chambers each containing an
electrochemical cell. These cells are based on 3 individually
addressed on-chip electrodes: working (WE; disk-shaped, 3 mm
diameter) and counter (CE) gold electrodes, and a reference silver/
silver chloride electrode (RE). The reaction chambers are defined in
photoresist, resulting in round-shaped 9 mm diameter chambers
with a sidewall height of 22 mm. The photoresist also passivates
connecting wires on the chip from the tested solution.
The device is microfabricated on a silicon wafer (single-side
polished prime grade quality, Ultrasil) modified with 1 mm of SiO2
passivation layer, grown with a plasma-enhanced chemical vapor
deposition (PECVD, Oxford Instruments) tool. A 20 nm thick Cr
layer followed by a 200 nm thick Au layer are deposited (DC
sputtering, AJA International, Inc.) on the wafer. The configuration
of the WE, CE, and RE is patterned in Au using photolithography
and thin film deposition technology, followed by subsequent
patterning of the reaction chambers in SU-8 2015 photoresist
(Microchem). Prior to the fabrication of the Ag/AgCl RE, an O2
plasma step (IPC 4000, Branson) is applied to clean the surface of
the electrodes from unwanted organic residues. The on-chip open
Ag/AgCl RE is fabricated by a 2-step electrodeposition method: 1)
Ag electroplating on the patterned Au electrode, 2) AgCl generation
on the Ag electrode surface [71].
2.3. Lab-on-a-chip electrochemical validation
The unmodified LOC was evaluated by testing its electrochemical performance for the analysis of a known redox reaction. The
ferrocyanide/ferricyanide reversible redox couple was used as a
model system to characterize the Nernstian electrochemical
response of the device. This was assayed by a conventional cyclic
voltammetry (CV) technique using 0.02, 0.05, 0.08, 0.1, and
0.15 V s1 scan rates (potential range between 0.4 V and 0.1 V;
2 cycles).
2.4. Catechol-chitosan redox cycling system fabrication, optimization,
and characterization
The catechol-chitosan redox cycling film fabrication was
modified from a recently published protocol [53]. Prior to the
catechol-chitosan fabrication steps, the electrochemical activity of
the device was verified with a ferrocyanide/ferricyanide solution.
The open circuit potential (OCP) and the cyclic voltammogram
response for each reaction chamber were recorded and compared
to the expected standard reduction potential and reversible
Nernstian behavior [72]. The Au WEs from all 4 reaction chambers
(Fig. 1A) were simultaneously coated with chitosan by shorting of
their contact pads and immersing the chip in chitosan solution
while applying 6 A m2 cathodic current. Platinum foil was used as
both the counter and the pseudo-reference electrodes. The
electrodeposition current was tested for an applied time of 15,
45, 60, and 90 seconds. This step resulted in a selective coating of
chitosan on the WE regions of the chip (Fig. 1B and 1C). Despite of
further rinsing step with PBS, some regions of the device other
than the WE (e.g., the CE) were left with residues of chitosan
(Fig. S1. CE’s averaged roughness increased from 3.7 to 15.5 nm
following the chitosan electrodeposition step). These residues can
contribute to additional variability in the electrochemical measurement across chambers. The device was mounted onto an
Fig. 1. (A) Image of the fabricated multi-chamber redox cycling integrated LOC (3 cm 4.5 cm). (B) Microscope image and (C) cross-section thickness profile of a chitosanmodified WE and the CE (black dashed line indicates the location of the scanned profile).
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
electronic interface platform and connected to a multiplexer and a
potentiostat. The chitosan films were further modified by placing a
5 mmol dm3 catechol solution in each chamber and applying 0.6 V
anodic potential to electrochemically oxidize the catechol.
Catechol oxidation was applied for a duration of 30, 60, 120,
180, 210, 240, and 300 seconds. To clean the chip from unbound
catechol, the chip was left in DI water for 5 minutes. The thickness
of the chitosan-only and catechol-chitosan films was characterized
using a surface profilometer (Dektak 6M, Veeco) followed by
drying of the film with a N2 gun.
2.5. Clozapine electrochemical detection assays
The multi-chamber device with unmodified WEs was used to
characterize the influence of Ru on the CLZ measurement. PBS
solutions with and without Ru were spiked with 0.5 (0.16), 1 (0.33),
5 (1.63), 10 (3.27), and 25 (8.17) mmol dm3 (mg mL1) CLZ
concentrations, and tested using a CV technique (potential range
between 0.4 V and 0.7 V; a scan rate of 0.02 V s1; 2 cycles). Devices
modified with different chitosan thicknesses and oxidized catechol
densities were tested to optimize the fabrication steps of the
catechol-chitosan film. Prior to each experiment, an initialization
step was performed to reduce most of the catechol embedded on the
chitosan film, charging the redox cycling system. Specifically, a CV
technique (potential range between 0.4 V and 0.6 V; a scan rate of
0.02 V s1; 2 cycles) was used to cycle redox reactions with the
initialization solution containing Ru as the catechol-reducing
mediator and Fc as the catechol-oxidizing mediator. The sensing
performance of CLZ was evaluated by testing a PBS solution
containing Ru and CLZ with a CV technique (potential range between
0.4 V and 0.6 V; a scan rate of 0.02 V s1; 1 cycle).
The sensing performance of the device was compared between
the bare and chitosan-modified Au WEs. Two chambers were
modified with catechol-chitosan films and two chambers were
modified with chitosan only. PBS solutions containing Ru and 0.16,
0.33, 1.63, 3.27, and 8.17 mg mL1 CLZ concentrations were tested
using CV technique (potential range between 0.4 V and 0.7 V; a
scan rate of 0.02 V s1; 2 cycles) or differential pulse voltammetry
(DPV) technique (potential sweep between 0 V and 0.7 V; potential
increment of 0.001 V; amplitude of 0.05 V; pulse width of 0.2 s;
sample width of 0.0167 seconds; pulse period of 0.5 s). Prior to each
DPV test, the catechol-chitosan film was charged in the presence of
Ru by applying a constant potential of 0.4 V for 120 s. Another set
of partially modified devices (i.e., 2 chambers were modified with
catechol-chitosan film and 2 chambers were modified with
chitosan only) were used to test CLZ sensing in human serum.
Commercial undiluted human serum was spiked with Ru and 0.33,
1.63, and 3.27 mg mL1 CLZ concentrations right before testing. The
CLZ-spiked serum samples were tested using a CV technique
(potential range between 0.4 V and 0.6 V; a scan rate of 0.02 V s1;
1 cycle). All electrochemical tests were done in duplicates. Limit of
Detection (LOD) values were determined as three standard
deviation units above the calculated Y-intercept from a linear
regression line [73].
3. Results and Discussion
3.1. Electrochemical characterization of the lab-on-a-chip with
unmodified working electrodes
3.1.1. Electrochemical performance validation
The electrochemical performance of the LOC was validated with
the commonly used redox couple ferrocyanide/ferricyanide. The
resulting cyclic voltammograms represented reversible Nernstian
characteristics for all the chambers in the device (Fig. 2A).
Repeatable testing showed a steady open circuit potential (OCP)
of 0.16600 0.00035 V, representing a stable on-chip open
reference electrode. Fig. 2B demonstrates a linear relationship
between the current peaks and the square root of the scan rate,
with slope values of 50 105 4 105 A s1/2 V1/2 and 49.0
105 1.5 105 A s1/2 V1/2, intercept values of 16 106 7
106 A and 20 106 3 106 A and R2 values of 0.98 and
0.99, for the anodic and cathodic current peaks respectively.
As steady-state conditions cannot be expected during the
electrochemical measurements in the LOC (steady state is true
when v << 8 106 V s1 [72]), currents that are resulted by linear
Fig. 2. Electrochemical validation of the unmodified LOC with a ferrocyanide/ferricyanide redox couple. (A) Representative cyclic voltammograms at scan rates of 0.02, 0.05,
0.08, 0.1, and 0.15 V s1. (B) The effect of the square root of the scan rate on the anodic and the cathodic peak currents. (C) The influence of the scan rate on the anodic and
cathodic peak potentials.
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
diffusion will dominate those resulted by radial diffusion.
Therefore it is valid to assume a theoretical relationship between
the peak current and the scan rate described by the Randles-Sevcik
equation [72] in order to calculate electrode active surface area
values, as follows:
n3=2 AD1=2 C v1=2
Ip ¼ 0:4463 F 3 =RT
Where Ip is the peak current [A], F [C mol ] is the Faraday constant,
R [J mol1 K1] is the universal gas constant, T [K] is the absolute
temperature, n is the number of moles of electrons transferred in
the cell reaction (n = 1 for a ferrocyanide/ferricyanide redox
reaction), A is the active surface area of the electrode [cm2], D
[cm2 s1] is the diffusion coefficient of the electro-active species
(0.72 105 cm2 s1 for ferricyanide and 0.67 105 cm2 s1 for
ferrocyanide [74]), C* [mol cm3] is the bulk concentration of the
electro-active species, and v [V s1] is the linear potential scan rate.
The averaged value of the active surface area of the electrode
calculated from both the anodic and cathodic slopes was A = 0.14
0.01 cm2. This value is approximately 2-fold bigger than the
calculated area of the disk electrode. This variation can be mainly
attributed to the additional active surface area of the electrode's
wire connector that is not covered by the SU-8 passivation layer.
Fig. 2C shows a positive relationship between the scan rate and
the anodic and the cathodic peak potentials, demonstrating
increasing peak separation for larger scan rates. The trends for
both reactions suggest that the sweep of the potential is not linear
anymore due to an increased potential drop at the interface of the
electrode. Such potential drop is resulted by uncompensated
resistance and may be dominant at a microliter volume solution.
Higher currents that are generated at faster scan rates increase this
potential drop, causing the peak potential to be a function of the
scan rate. As these limitations are common and can be
compensated in micro-scale electrochemical devices, the electrochemical performance of the device was validated.
3.1.2. Clozapine electrochemical oxidation
The effect of Ru on the electrochemical oxidation of CLZ was
tested with a bare gold WE in the LOC. Fig. 3 shows CVs of PBS
solutions containing 0, 3.27, and 8.17 mg mL1 CLZ, in the presence
and absence of Ru. When Ru was absent, increasing concentrationdependent anodic (Ep = 0.33 V) and cathodic (Ep = 0.27 V) currents
of CLZ were shown (Fig. 3A). When Ru was introduced to the
sample, a sharp anodic peak became apparent at 0.43 V (Fig. 3B).
This peak partially masked the CLZ anodic peak current. This
anodic reaction was also present when no CLZ was added to the
sample, suggesting that Ru may be involved in overlapping the
electrochemical reaction of CLZ.
These results are significantly different from the previous study
with buffered solution containing lower chloride ions concentrations [53], which did not show this peak in the presence of Ru,
suggesting that chloride ions are also involved in the reaction. A
possible explanation can be an electrochemical reaction between Ru
and chloride ions. For example, solid ruthenium (reduced from
ruthenium(II); E0 = 0.23 V) can be oxidized in the presence of
chloride ions resulting in ruthenium trichloride at an E0 = 0.46 V. In
the current LOC scheme, it is necessary to perform the test in the
presence of chloride ions to maintain a stable half cell potential for
the on-chip open Ag/AgCl reference electrode. Additionally, the
presence of chloride ions better simulates the levels present in
normal human serum, ranging between 100 to 110 mmol dm3 of
chloride ions [75,76]. Identifying the source of electrochemical
reactions that occur at potentials adjacent to the CLZ standard
reduction potential will help isolating the signal from CLZ oxidation.
3.2. Catechol-chitosan redox cycling system integration
3.2.1. Clozapine anodic current amplification
The ability of the catechol-chitosan modified LOC to amplify
CLZ signal was studied. Fig. 3B (8.17 mg mL1 CLZ concentration
curve) and Fig. 4 show CVs of CLZ measured with bare electrodes
Fig. 3. CVs of 0 mg mL1 (solid black), 3.26 mg mL1 (dot-dashed blue), and 8.17 mg mL1 (dashed magenta) CLZ concentrations in the absence (A) and the presence (B) of Ru in
PBS solutions. Right-hand graphs present inside view of the CLZ anodic peak current density range. (For interpretation of the references to color in this figure legend, the
reader is referred to the web version of this article.)
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
Fig. 4. Cyclic voltammograms of PBS solution with and without CLZ measured with electrodes modified with either catechol-chitosan (without CLZ – dashed green; with CLZ
– dot-dot-dashed magenta) or chitosan only (without CLZ – solid black; with CLZ – dot-dashed blue). Magnified graph presents the oxidative region of the cyclic
voltammogram. CLZ concentration was 8.17 mg mL1. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
and electrodes modified with catechol-chitosan and chitosan films,
respectively. The overall anodic currents measured with catecholchitosan modified electrodes were larger than their counterparts
measured with either bare or chitosan-modified electrodes. These
amplified currents are due to the redox cycling system, which
continuously regenerates CLZ and other oxidized and reversible
electro-active species in the sample. Despite the amplified current
of CLZ, its anodic peak wasn't differentiable in the current setup
from other anodic reactions in the system. Therefore, different data
processing approaches were evaluated and assessed with respect
to their effect on the sensing performance.
Evaluating the electrochemical current generated at the
expected oxidation potential of CLZ provides a specific approach
as opposed to calculating the total electrochemical charge for a
range of potentials. On the other hand, the calculated charge
essentially integrates a set of CLZ oxidative currents at different
potentials, accumulates the whole CLZ oxidation capacity in the
system, and eventually obtains higher signals. However, during
this process, oxidation capacities of other electro-active species
can be accumulated as well, which may contribute to higher
background signals. Such background signals can be quantified and
subtracted by the analysis of control samples (e.g., negative control
without CLZ) and transducers (e.g., bare or chitosan-only modified
sensor). The electrochemical charge was calculated for 4 different
sections of the CV signal (Fig. 5) to evaluate which region better
correlates to CLZ sensing: 1) Qr_red for cathodic charge at E < 0 V,
2) Qr_ox for anodic charge at E < 0 V, 3) Qo_ox for anodic charge at
E > 0 V, and 4) Qo_red for cathodic charge at E > 0 V. Two different
data subtraction methods were used to process the charge values
(Eqs. (2) and (3)):
“Sequential testing” approach : Q i ¼ Q i; with CLZ Q i;without CLZ electrode type
magnitudes containing lower percentage errors with the “parallel
testing” approach compared to “sequential testing” approach.
Furthermore, Qr_red and Qo_ox yielded a higher magnitude of
values than those processed using Qr_ox and Qo_red for both
“parallel testing” and “sequential testing” approaches due to CLZ
oxidation amplification and the related Ru reduction dominant in
these sections. Within the “parallel testing” approach, using either
the charge calculated from the bare or the chitosan-modified
electrodes didn't have a significant effect on the values and errors
calculated for Q*o_ox (50 mC and 10% vs. 45 mC and 13%). For
Q*r_red values, the chitosan only electrode is suggested to be a
better reference than the bare electrode by demonstrating higher
anodic values and better percentage error (90 mC and 17% vs.
60 mC and 27%). This analysis shows that “parallel testing”
approach using Q*o_ox is better to identify CLZ detection.
Aside from analyzing data processing approaches, the methodological and design aspects of a POC testing device have to be
taken into consideration. To provide accurate reference values for a
“sequential testing” approach, two different solutions (i.e. a
reference solution without CLZ followed by the sample with
CLZ) will have to be introduced to the detection chamber and to be
“Parallel testing” approach : Q i ¼ Q i;catecholchitosan Q i; ref erence
with CLZ
where the annotation i is the relevant CV section (r_red; r_ox; o_ox;
o_red), “electrode type” refers to bare, chitosan only and catecholchitosan modified electrodes, and “reference” denotes the
electrode used to measure the reference signal (either bare or
chitosan only electrodes).
Table 1 compares the processed charge values for the two
different data subtraction approaches. Results show higher charge
Fig. 5. Cyclic voltammogram charge sectioning diagram; Qr_red (red-yellow) for
cathodic charge at E < 0 V (Ru reduction), Qr_ox (blue) for anodic charge at E < 0 V,
Qo_ox (purple) for anodic charge at E > 0 V (CLZ oxidation), and Qo_red (green) for
cathodic charge at E > 0 V. (For interpretation of the references to color in this figure
legend, the reader is referred to the web version of this article.)
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
Table 1
Comparison of the different charge values between the 2 different data subtraction approaches.
Data subtraction approach
“Sequential testing”
“Parallel testing”
Electrode type
chitosan only
Reference = bare
Reference = chitosan only
Percentage error/%
Percentage error/%
Percentage error/%
Percentage error/%
13.0 1.3
0.005 0.020
5.0 1.0
0.003 0.100
4.0 2.6
0.06 0.24
1.4 0.3
3 24
0.9 2.0
9 15
2.0 0.6
60 16
6.0 1.5
50 5
5.0 0.3
90 15
4.0 1.5
45 6
3.6 0.3
measured by the same electrode. For a “parallel testing” approach,
only one solution (i.e. the tested sample with CLZ) will have to be
introduced to the chamber while a differential test will be
performed between the catechol-chitosan and chitosan only
modified electrodes. Previous experience with the catecholchitosan film showed that even though the system is highly
repeatable in buffered solutions, introduction of human serum
samples in consecutive tests resulted in fouling of the sensor with
approximately 10% decrease of the signal [53]. From the results in
Table 1 and the methodological and design considerations, the
“parallel testing” approach would provide a sensitive LOC with
minimal pretreatment steps. As both bare and chitosan only
modified electrodes showed similar detection factor values,
chitosan was chosen to be the reference signal. Utilizing electrodes
coated with chitosan film provides possible future chemical
modifications (e.g. moieties that are negatively charged to repulse
CLZ or that can adsorb CLZ decreasing its diffusion coefficient).
These modifications will interact with CLZ, selectively decreasing
the reference electro-active charge generated by CLZ.
A comparison of the calculated “parallel testing” approach value
for Q*o_ox (defined as CLZ detection factor) between macro-scale
electrodes [53] and the LOC presented here showed higher values
for the former (263 mC vs. 45 mC, respectively). However, when
taking the surface area of the electrode into account, the charge
density values were similar (468 mC cm2 vs. 450 mC cm2).
Therefore, in spite of the macro-scale setup having electrodes
that have approximately 5 times higher surface area and millilitersized solution volumes, the LOC device achieved the same CLZ
detection factor with microliter-sized volume samples and in the
presence of chloride ions.
3.2.2. Catechol-chitosan film optimization
The chitosan electrodeposition and catechol modification steps
were optimized to maximize the ability of the LOC to detect CLZ. First,
the influence of the electrodeposition duration on the resulting
chitosan thickness was characterized (Fig. 6A). Results showed a
positive linear relationship with a slope value of 14.00 0.25 nm s1,
an intercept value of 180 15 nm and R2 value of 0.99. Because the
thickness values include the dried chitosan layer as well as the
underlying Cr/Au electrode layers, the intercept value indicates the
thickness of a bare electrode. The intercept value was smaller than
expected (220 nm). This variation may be due remaining chitosan
residues on the SiO2 surface adjacent to the electrode. The linear
relationship between chitosan thickness and electrodeposition step
duration suggested good process control in the micro-chip level with
expected deposition mechanism [77].
Fig. 6. Optimization of redox cycling film preparation. (A) The influence of the chitosan electrodeposition time on the thickness of the electrodeposited chitosan layer.
(B) Oxidized catechol amount for increasing oxidation step duration. (C) The effect of the chitosan electrodeposition and the catechol modification steps on the CLZ detection
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
Fig. 7. CLZ dose response characteristics. (A) A comparison of the LOC integrated with the catechol-chitosan modification (black squares) and the bare gold electrode (red
circles). (B) A comparison of electrochemical techniques measured with the LOC integrated with the catechol-chitosan modification: CV (black squares) and DPV (red circles).
Negative values were observed for bare electrode due to higher chitosan reference charge values. (For interpretation of the references to color in this figure legend, the reader
is referred to the web version of this article.)
3.3. Clozapine sensing performance with the catechol-chitosan redox
cycling integrated lab-on-a-chip
The sensing performance of the LOC was characterized using
different concentrations of CLZ (0.16, 0.33, 1.63, 3.27, and 8.17 mg
mL1) in PBS solutions. Fig. 7A compares differently modified
electrodes integrated in the LOC for the CLZ calibration curve
obtained from CV results (Corresponding anodic current density
peak values are presented in figure S2). Positive linear relationships were observed for the modified and the bare electrodes, with
sensitivity values of 54 7 mC mL cm2mg1 and 7.0 2.6 mC
mL cm2mg1, and LOD values of 0.8 mg mL1 and 2.1 mg mL1,
respectively (Corresponding analysis of the anodic current density
peak values of the modified electrode resulted in similar LOD
values but inferior sensitivity; Fig. S3 and Table S1). The better
sensitivity (8-fold higher) and better LOD (2-fold lower) measured
with the catechol-chitosan modification compared to bare electrodes are related to the higher charge values achieved due to redox
cycling. These higher charge values increase the signal-to-noise
ratio and the selectivity of the sensor for CLZ.
DPV is an electrochemical technique that is commonly used in
analytical applications due to decreased non-faradaic currents
during measurements. Fig. 7B presents the effect of electrochemical techniques on the sensing performance of the LOC integrated
with the catechol-chitosan film (Corresponding anodic current
density peak values are presented in Fig. S2). A positive linear dose
response for the DPV technique was obtained with sensitivity and
LOD values of 0.100 0.013 mC mL cm2 mg1 and 0.9 mg mL1,
respectively. CV demonstrated better sensitivity (540-fold higher)
than DPV even though higher non-faradaic currents were
measured. As DPV is a preferred technique for electrochemical
reaction analysis, CV will be a better choice for sensing purposes
due to the higher accumulated charge. On the other hand, as Ru
seems to interfere with CLZ detection, DPV was used here to test
CLZ in the absence of Ru as opposed to CV. Negative control
measurement showed a small anodic current peak at 0.34 V which
may be due to Ru interference (Fig. S4). This suggests that Ru
compounds may remain in the catechol-chitosan film due to the
prior electrochemical charging step. Integration of methods to
ensure that Ru does not remain in the film (e.g., soaking in DI for Ru
diffusion out of the film) will decrease potential interference,
reduce background signal, and improve the sensing performance.
3.4. Clozapine detection in human serum
The ability of the redox cycling integrated LOC to detect CLZ in
human serum with no sample preparation steps (dilution,
filtration, etc.) was tested. Different concentrations of CLZ (0.33,
1.63 and 3.27 mg mL1) spiked into undiluted serum were analyzed
with the device. The intensity of electrochemical current generated
from commercial serum varied by a magnitude of order between
batches (data not shown) [78]. Therefore, an additional pair of
catechol-chitosan and chitosan- modified electrodes were used as
an on-chip reference measurement subtracting the blank signal
from serum with no consecutive tests that may foul the sensor
(Eq. (4)):
Q oo x;
serum with CLZ
Q*o_ox w/ CLZ - Q*o_ox w/o CLZ / µC cm -2
CLZ detection was evaluated for various durations of chitosan
electrodeposition (15, 45, 60, and 90 seconds) and catechol
modification (30, 60, 120, 180, 210, 240, and 300 seconds).
Fig. 6C shows a 2D contour representation of CLZ detection factor
values tested with various chitosan thicknesses and oxidized
catechol density (calculated from the generated charge during the
catechol electrochemical oxidation step; Fig. 6B). An overall trend
of increasing CLZ detection factor was demonstrated for increasing
oxidized catechol densities. An optimized region was observed
with a chitosan thickness of 800 nm and an oxidized catechol
amount of 34 nmol (converted from the charge generated during
the catechol oxidation step using Faraday’s law), corresponding to
a chitosan electrodeposition time of 60 seconds and a catechol
modification time of 240 seconds, respectively.
Q oo x; serum only
CLZ concentration / µg mL
Fig. 8. Background subtracted detection factor values of human serum spiked with
either 0.33, 1.63, or 3.27 mg mL1 CLZ concentrations.
H. Ben-Yoav et al. / Electrochimica Acta 163 (2015) 260–270
Fig. 8 shows the resulting subtracted values for different
concentrations of CLZ-spiked serum. These results show increasing
charge values with higher CLZ concentrations. A statistical analysis
using an independent two-sample t-test demonstrated a significant difference between 0.33 and 3.27 mg mL1 CLZ concentrations
with a p-value of 0.037 (significance level of 0.05). These results
provided the feasibility of the device to differentiate between
different concentrations of CLZ in untreated human serum
samples. As these concentrations are close to the upper limit of
the required clinical range for toxicity evaluation (i.e., 1 mg mL1),
additional comprehensive characterization of the capability of the
device to detect CLZ in serum will provide the required sensitivity
and LOD values. For example, the variability of commercial serum
samples will be addressed by acquiring samples directly from
healthy subjects and characterizing the subject-to-subject electrochemical activity difference. Furthermore, the cross-reactivity
of different electro-active species will be investigated to identify
the major interferents.
include possible variability and cross-reactivity parameters
(e.g., electro-active species, patient-to-patient variability) interfering with the capability of the device to detect electrochemical
signals generated by CLZ. Another direction will investigate other
redox cycling schemes instead of the catechol-chitosan film, such
as a 2-electrode configuration. As such a scheme will decrease the
complexity of the fabricated device, other aspects will have to be
taken into consideration during the design step (e.g., diffusion
limitations due to the distance between electrodes, unstable half
cell potential of the electrodes due to variations of the electoractive species concentrations). This novel application of portable
analytical micro-systems for schizophrenia treatment monitoring
will provide advantageous characteristics such as rapid, on-site,
cost-effective, and straight-forward testing of the efficacy and the
safety of the treatment at the POC (e.g., pharmacy, physician's
office, clinic, hospital or at-home). With these devices the cost and
burden of monitoring could be reduced, increasing the compliance
of CLZ treatment among patients and prescribers and revolutionize
the paradigm of how mental disorders are currently managed.
4. Conclusions
In this work, we present a redox cycling integrated LOC for real
time, electrochemical detection of CLZ. The microfabricated
electrochemical device demonstrated highly repeatable and
Nernstian behavior in multiple reaction chambers with microliter
sample aliquots. Different electrochemical techniques, data
processing approaches, and redox cycling film fabrication parameters were evaluated in order to optimize CLZ sensing and signal
amplification. Studying the influence of the catechol-chitosan film
fabrication process on physical and chemical reactions participating in the sensing mechanism will elucidate dominant phenomena
that can be used to improve drug accessibility and limit ionic
resistance in the system. The sensitivity and the LOD of the redox
cycling integrated device were studied with buffered samples,
providing crucial evidence of the sensor functionality. Finally, the
feasibility of the device to differentiate between two concentrations of CLZ in human serum within the upper limit of the
required clinical range was demonstrated.
These achieved characteristics demonstrate CLZ detection with
a miniaturized redox cycling system for the first time. While
macro-scale assays showed an 8-fold LOD improvement [53] and
the recommended therapeutic CLZ blood levels are in the range of
0.35–1 mg mL1 [13], these accomplishments provide initial
understanding of systems integration and dominant mechanisms
in redox cycling integrated electrochemical micro-systems for
analytical applications. Other redox cycling-integrated miniaturized detection assays demonstrated similar characteristics for
other analytes [79–82]; for example, 10 mmol dm3 LOD of
paracetamol was demonstrated by Goluch et al. [79]. Despite
the better LOD and sensitivity of current commercial benchtop
technologies for CLZ detection (e.g., liquid chromatography–
tandem mass spectroscopy quantification method [83]), this
miniaturized analytical device will allow treatment teams to
perform analysis at the POC in a low cost, fast, and straightforward
way aiming to guide CLZ dosages within the effective range, to
decrease the patient's burden, and to personalize medical care.
This work is the first step in the device development and
characterization towards portable micro-systems for rapid analysis of CLZ with minimal pretreatment steps. Aiming towards an
assay with minimal pretreatment steps provides a simpler LOC
paradigm that will achieve better overall systems integration
capabilities. However, such a paradigm will suffer from limited
sensitivity and LOD as sample pretreatment steps decrease the
level of interference. A future study of CLZ detection in serum
samples will provide additional information on the device
performance with complex biological fluids. Such study will
The authors would like to thank the Robert W. Deutsch
Foundation, the Maryland Innovation Initiative (MII), and the NSF
Grant No. DGE0750616 for financial support. The authors likewise
appreciate the support of the Maryland NanoCenter and its FabLab.
The authors wish to thank Professor Yosi Shacham-Diamand from
Tel-Aviv University for providing experimental expertise in
on-chip reference electrode fabrication. The authors wish to also
thank the Bioengineering rotation students in MSAL: Bao-Ngoc
Nguyen, Bharath Ramaswamy, Stephan Restaino, and Nicholas
Woolsey and the undergraduate intern Gillian Costa for the useful
Appendix A. Supplementary data
Supplementary data associated with this article can be found, in
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