Chemotaxis-based smart drug delivery of epirubicin using a 3D printed mi-
crofluidic chip
Kolsoum Dalvand, A. Ghiasvand, Vipul Gupta, Brett Paull
PII: S1570-0232(20)31332-5
DOI: https://doi.org/10.1016/j.jchromb.2020.122456
Reference: CHROMB 122456
To appear in: Journal of Chromatography B
Received Date: 22 July 2020
Revised Date: 12 October 2020
Accepted Date: 16 November 2020
Please cite this article as: K. Dalvand, A. Ghiasvand, V. Gupta, B. Paull, Chemotaxis-based smart drug delivery of epirubicin using a 3D printed microfluidic chip, Journal of Chromatography B (2020), doi: https://doi.org/ 10.1016/j.jchromb.2020.122456
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Chemotaxis-based smart drug delivery of epirubicin using a 3D printed
microfluidic chip
Kolsoum Dalvand , A. Ghiasvand , Vipul Gupta , and Brett Paull
Department of Chemistry, Lorestan University, Khoramabad, Iran
Australian Centre for Research on Separation Science (ACROSS), School of Natural Sciences, University of
Tasmania, Hobart, Tasmania 7001, Australia
Alireza Ghiasvand, E-mail: [email protected]
Tel-Fax: +61 3 6226 2121
ORCID ID: https://orcid.org/0000-0002-4570-7988
Highlights:
Drug delivery performed via chemotaxis in a 3D printed microfluidic device.
A self-propelled ionic liquid drop was used for targeted delivery of epirubicin.
A central composite design was used to optimize the drug loading variables.
Abstract
Recent developments on self-propelled microdroplets, moving controllably in response to an external stimulus
like chemical, electrical, or magnetic field, have opened a new horizon for smart drug delivery investigations. On
1
the other hand, the new achievements in 3D printing technology has provided a promising option for the
fabrication of microfluidic devices, which is an unrivalled platform for in-vitro drug delivery studies. By
synergizing the features of chemotaxis, 3D printing, and microfluidic techniques a new approach was introduced
to deliver the drug to targeted sites with a well-controlled method and a reasonable speed. A self-propelled ionic liquid ([P6,6,6,14][Cl]) microdroplet, as the drug carrier, was utilised for the targeted delivery of epirubicin
anticancer drug within an integrated drug delivery microfluidic system. The asymmetric diffusion of [P
6,6,6,14
] ion
from the microdroplet into an aqueous solution with chloride gradient concentration (created under an external
electrical field) caused the microdroplet to move. The spatial and temporal position of the moving microdroplet
could be controlled by changing the magnitude and polarity of the external electrical field. A piece of hollow-
fiber, fixed next to the anode, was filled with phosphate buffer (as the receptor) and used to remove the drug from
the carrier. The receptor solution was then taken and injected into a HPLC system for quantification of the released
drug. After one-at-a-time optimization of the channel geometry and electrolyte concentration, the experimental
variables affecting the drug loading including contact time, pH, and volume of carrier were optimized via a central composite design (CCD) approach.
Keywords: Smart drug delivery; Chemotaxis; 3D printing; Microfluidics
1. Introduction
Chemotaxis is the ability of movement of cells and living organisms in response to an external stimulus, present
within their surrounding environment. It has been observed in a wide variety of microorganisms, plants, and
animals, as a vital process for the living creatures [1]. Chemotaxis has been extensively studied during two last
decades due to its fundamental role in cancer metastasis, embryogenesis, cell development, inflammation, and
wound healing [2, 3]. It is driving force for the movement of sperm towards ovum during mammalian fertilisation.
Because of its momentous effect on the cell migration and biological processes, many studies have been conducted
to understand the mechanism of chemotaxis [4]. For this purpose, different techniques including swimming plate,
capillary assay, agar plug assay, PP-technique, Boyden’s chamber, Zigmond chamber, and Dunn’s method have
been developed [5, 6]. Although, these techniques mostly provide a qualitative approach to measure chemotaxis
response, but they have low reproducibility and sensitivity and usually take several hours to several days to be
accomplished. Different microfluidic setups have been developed to produce, control, identify, and quantify
chemotaxis of cells of different types [7, 8]. Compared with other relevant techniques, microfluidic-based
chemotactic systems can generate accurately controlled gradients and enable dynamic observation and
quantitative study for bacterial chemotaxis [9]. Chemotactic-driven phenomenon has inspired scientists to construct synthetic microdroplets, which are able to exhibit similar behavior in-vitro.
Different chemotactic stimulating mechanisms including light [10], chemical [11], electrochemical [12], magnetic
field [13], electrical field [14], temperature gradients [15], and acoustic waves [16] have been reported to actuate
the microdroplets and stimulate them to move. Electrotaxis or galvanotaxis is the movement of cells in response
to an electric field and has better control over the direction and speed of cell migration compared to chemotaxis.
Electrotaxis has been applied in both modes of in-vivo and in-vitro [17, 18]. To study the effect of electrical
stimulation on immune cell activation during wound healing an in-vivo electrotaxis system was developed by
Wang et al. [19]. They investigated the effect of electrical stimulation in the rat to monitor inflammation in a wound. In another study, Feng et al. studied the electrotactic-guided cells in a rat brain model [20].
Cancer is now a primary cause of death and a major health burden in many societies. To control this so-called big
worldwide killer different anticancer and chemotherapy drugs have been developed [21]. Most of the
chemotherapy medications do not have selective effects against cancer cells and can also kill normal cells.
Therefore, a big challenge in cancer treatment is to deliver the drug directly to the tumour, without damaging the
healthy tissues [22, 23]. Different efforts have been devoted to address this issue like “smart drug delivery” by
using nano/submicron droplets [24]. Smart drug delivery investigations struggle to find proper pathways to deliver
the drug to the right area, at the right time, and with a controlled dose. This technology has raised many hopes to
improve the treatment of incurable diseases like cancer [25]. Among different stimulus-based drug delivery
strategies, electrotactic driven ionic liquid (IL) microdroplets have been noted more for their simplicity, versatility, reliability, and proper reproducibility in drug release [26].
ILs are organic salts that are liquid at low temperatures. Good thermal stability, low combustibility, good solvation
ability, and very low vapor pressure have made ILs ideal solvents for targeted drug delivery [27].
microdroplets in aqueous solutions. It also benefits a wide electrochemical redox window, which makes it stable
in electrolytic systems across a broad potential range. In this regard, Moniruzzaman et al. reported an IL-in-oil
microemulsion system for the delivery of very low soluble drugs to enhance their potential for topical and transdermal treatment [28].
Epirubicin is one of the most used chemotherapy drugs that its toxicity and low selectivity significantly reduces
the efficacy and therapeutic effect of this antibiotic. Therefore, the development of alternative methods for targeted
delivery of epirubicin and other anticancer drugs to tumour cells is of crucial importance. Malaekeh-Nikouei et
al. used mesoporous silica nanoparticles as carriers for epirubicin hydrochloride through in-vitro and in-vivo
methods to improve the antitumor efficacy of epirubicin [29]. They found that the drug release pattern was pH-
and time-dependent. Similarly, Deng et al. loaded epirubicin on polysialic acid/polyethylene glycol conjugate-
modified liposomes to enhance its anticancer performance [30]. Despite the attributed features, these drug delivery
systems are inactive and the drug-loaded micro- and nano-carriers lack the driving force for efficient delivery of
drug to the tissue. Alternatively, the use of active drug delivery systems that are biocompatible with their function
to swim and penetrate the tissues for drug release at the right time and place have been the subject of much
researches, recently. These systems can be controlled using an external source like electrical field, magnetic field, light, and ultrasonic waves [31].
This research aimed to investigate smart delivery and release of epirubicin using a drug-loaded IL microdroplet,
controlled reversibly by a low-voltage electrical field. A microfluidic device was 3D printed and equipped with a
proper receptor at the end of the channel to mimic a living tissue. The receptor was a hollow-fiber filled with a
few microliters of phosphate buffer and removed the drug from the microdroplet gradually. The receiving buffer
was then removed using a microsyringe and injected into a HPLC system to quantify the released drug. The
affecting experimental variables were optimized through a response surface method (RSM), based on central composite design (CCD).
2. Materials and methods
2.1. Chemicals and solutions
Trihexyltetradecylphosphonium chloride ([P6,6,6,14][Cl]) was provided by Sigma-Aldrich (St Louis, MO, USA).
Sodium chloride (NaCl), dichloromethane, acetone, disodium hydrogen phosphate, potassium dihydrogen
phosphate, ammonium dihydrogen phosphate, and phosphoric acid (85%) were purchased from Merck
(Darmstadt, Germany). Epirubicin hydrochloride (EP) was purchased from EBEWE Pharma GmbH Nfg KG
(Unterach, Austria). Liquid chromatographic grade methanol was sourced from Merck. Stock standard solution of EP was prepared in deionized water (5 mg L ) and stored in a fridge at 4 °C in the dark.
2.2. Instruments and HPLC conditions
A Shimadzu HPLC system (Kyoto, Japan) consisted of two reciprocating pumps, an SPD-10AD ultraviolet
detector, a DGU-14A degasser, and a high-pressure injection valve (with a 20-μL injection loop) was used for
4
chromatographic analysis. A Wakosil II 5C18 RS column (250 mm × 4.6 mm × 5 μm diameter particles) was
used at constant temperature of 35 °C. An isocratic solvent program using a mixture of phosphate buffer (0.05 M)
and methanol (40:60) at a flow rate of 1.0 mL min was used. The buffer was first adjusted to pH 3.0 and then
filtered through a 0.45 µm pore size filter membrane (Millipore). The detector was set at 254 nm. A TP120-10S
DC power supply was used to apply voltage. A Shimadzu UV-1650 PC spectrophotometer was used to record absorption spectra of epirubicin.
2.3. Fabrication of the mmicrofluidic chip
The microfluidic devices with different channel geometries (triangular, rectangular, and half-circle) and lengths
were 3D printed using a PolyJet printer. The models were designed using Fusion 360 (Version 2018, Autodesk
Inc., San Rafael, CA, USA) and printed using a professional Objet Eden260VS 3D printer from Stratasys (Eden
Prairie, MN, USA) with a resolution of 600 × 600 × 1600 DPI. All parts of the model’s body were made from a
single build material and the void spaces filled with the support material, which was washed out after completion
of the fabrication process. VeroClear-RGD810 and SUP707 from MatterHackers (Foothill Ranch, CA, USA),
were used as the build material and the water-soluble support, respectively. A Stuart Orbital Platform Shaker SO1
(Stuart Scientific, Staffordshire, UK) was used for washing and removing the support material from the
microfluidic devices. The printer can print 60 such chips in 250 min, resulting in ca. 4 min of print time per chip.
Immediately after printing, the surplus layers and parts of the support were detached from the microfluidic devices
and they were soaked in 2% NaOH solution and rinsed two times. Then, they were socked in 2% NaOH solution
and shaken for 24 h using the rotary shaker, for complete dissolution of the support material and opening the
channels and holes. Finally, the microfluidic devices were dried at room temperature and stored in plastic bags. The scheme of the 3DP microfluidic device is shown in Fig. 1.
To obtain the optimized geometry and dimensions, different types of channels with various geometries were
fabricated. The body of the microfluidic devices were made from transparent material for clear photography and
video recording of the microdroplet motion. Two screw holes were embedded at both ends of the channel to
accommodate the electrodes. Two copper electrodes were fixed into proper PEEK finger-tight fittings and inserted into the electrode sinks.
2.4. Loading of the drug into the ionic liquid
5
Drug loading behavior of [P6,6,6,14][Cl] was preliminarily investigated using UV-Vis technique. For this purpose,
50 µL of the IL was added to a 500 µL aqueous solution of epirubicin hydrochloride (5 µg mL ) and shaken for
one min, to dissolve EP in the IL and establish an equilibration. It was then left stagnant for complete phase
separation. Finally, the aqueous phase was separated using a microsyringe, transferred into a microvolume cuvette,
and its EP concentration was measured by UV-Vis spectrophotometry. The top phase (IL) was used for the
chemotaxis investigations. The UV-Vis study showed that a 20-µL microdrop of the IL the can uptake 1.925 µg
of EP (96.4 µg mL ). The simple planer structure of epirubicin and trihexyltetradecylphosphonium ion are shown in Fig. 2.
2.5. Migration and chemotaxis assay in the microfluidic device
For chemotaxis experiments, the channel was filled with an electrolyte solution (200 µL, 0.01 M NaCl) and a 5 V
DC voltage was applied. After 20 seconds, a 20-µL IL microdroplet containing EP was placed into the channel at
the cathode side. The potential creates a concentration gradient of ions across the electrolyte solution by moving
Na ions toward the cathode and Cl ions toward the anode. Trihexyltetradecylphosphonium is as an effectual
ionic surfactant and when the IL microdroplet is placed in the channel, [P6,6,6,14
] cation diffuses out from the
microdroplet and disperses into the solution, through Marangoni effect.[32] Consequently, a sudden drop in the
surface tension is occurred and driven the microdroplet toward the anode. In other words, the potential-stimulated
asymmetric release of [P
6,6,6,14
] from the microdroplet leads to the electrotactic movement. The self-propelled IL
microdroplet moves along an aqueous-air boundary to the end of channel, inside the water molecules network.
The movement direction can be changed by changing the polarity of the electrodes and the speed can be controlled
by increase/decrease of voltage. The external electric field can also localise the negative and positive charges
inside the microdroplet and converts it into a bipolar electrode by Faradic rearrangement of the ions. This process
can make an ion concentration gradient that may amplify the electrotactic movement. Nevertheless, the effect of
bipolar phenomenon in the movement of microdroplet seems to be much smaller than the electrotactic mechanism, because of very strong ion-paring system inside the droplet [33].
3. Results and discussion
3.1. Influence of the microchannel geometry
To investigate the effect of geometry and dimensions of the microchannel, three types of channels (rectangular,
triangular, and half-circle geometry) with different dimensions were fabricated (Table 1). Each channel was filled
6
with the electrolyte solution and subjected to the electrotactic drug delivery experiment. Amount of the released
drug (EP), and the time required for the microdroplet to reach the end of the channel were recorded as the response.
The results demonstrated that amount of the released drug did not affected remarkably by the channel’s geometry
when the dimensions are the same. Nevertheless, the shortest movement time (time to cross the channel length)
was obtained for the half-circle channel. This can be attributed to the laminar flow that is provided by this type of channel. It is well known that laminar flow can exist in low Reynolds number (Re) as follow:
where U and μ are average velocity and viscosity of the fluid, and DH is hydraulic diameter, four times of the
cross-sectional area of the channel divided by the wetted perimeter [34]. The results showed that movement time
for a 20 × 2 × 1.5 mm half-circular channel was 20 s, while for rectangular and triangular channels with the same
dimensions it was 60 and 300 s, respectively. Therefore, half-circular channels were selected for further studies.
Furthermore, the internal channel dimensions were optimized. It was revealed that channels with shorter length
than 20 mm was not proper for visualisation and video recording of the movement process, while longer channels
led to microdroplet-splitting. Therefore, a 20 × 2 × 1.5 mm half-circular channel was chosen as the best choice for further investigations.
3.2. Electrolyte concentration and electrical current in the channel
Different concentrations of the electrolyte (NaCl) were examined for the electrotactic drug delivery and the current
generated between the electrodes was measured by varying voltage over the range of 1-12 V. The movement was
not satisfactory by using electrolyte concentrations lower than 0.001 M, even at high voltages. A concentration of
0.01 M led to a proper movement but caused electrolysis of the solution and generated bubbles in the cathode (at
the voltages more than 7 V), which interrupted the microdroplet movement. The bubble production appeared at
lower voltages than 7 V, when concentrations more than 0.01 M were used. Therefore, 0.001 M was considered
for further studies, which did not show bubbles at a wide range voltage up to 10 V and created proper microdroplet
movement with low current density. The results for 0.01 and 0.001 M concentrations are presented in Fig. 3. It
was proved that in addition to voltage and electrolyte concentration, the speed of the microdroplet depends on
other parameters like volume of the microdroplet, time to put the microdroplet in the channel, and contact area
between the microdroplet and the aqueous solution. If all these parameters are kept constant, velocity of the microdroplet increases with increasing the voltage (below the valuse that generate bubble).
7
3.3. Optimization of the drug loading conditions via CCD model
The most important variables that affect the drug loading process are pH of the EP solution (pH), contact time
between EP and the IL (Time), and volume of the carrier (Vc). These parameters were investigated using a
response surface methodology (RSM) based on a five-level CCD model. Table 2 shows the factors and their levels. The drug load percentage (R%) was calculated using the following equation (Eq. 2).
are the initial and equilibrium concentrations of EP in the aqueous solution. The predicted
experiments by the CCD model were carried out by considering R% as the response, as shown in Table 3. The
experimental results were evaluated by the analysis of variance (ANOVA) method. The results are summarized
in Table 4. As the results show, the p-values of the main variables were less than 0.05, while the p-values of the
Lack-of-Fit were higher than 0.05, confirming that all variables were statistically significant with 95% confidence. According to the ANOVA results, a quadratic model was fitted using the following equation:
R = 59.41 – 0.9184A – 4.25B + 6.15C + 0.2410AB + 0.2406AC – 4.36BC – 2.84A2 + 0.3911B2 – 2.13C2 (Eq. 3)
The F-value of Lack-of-Fit was 0.7666, which was not significant relative to the pure error and confirmed the
validity of the model [35]. Besides, predicted-R of 0.9522 showed that the regression model possessed a high
significance. R
2
and adjusted-R
were 0.9863 and 0.9739, respectively, indicating a desirable value for the model
validation. On the other hand, coefficient of variation (CV), sum of squares of the predicted residuals (PRESS),
and values of adequate precision (S/N) were 2.20, 53.50 and 1.24, respectively, which indicated the model
desirability and feasibility. Fig. 4 shows the response surface plot for drug load percentage as a function of two
variables, while the third variable is kept constant at its central level. Plotting three-dimensional graphs of response
surfaces are very effective in understanding the main effects and interactions of the variables. The nonlinear nature
of the three-dimensional response graphs indicates existence of significant interactions between each of the
independent variables and response. As the results show, volume of the carrier has a positive effect on increases
of response, while pH and time had negative effects. Overall, the results showed that the optimum conditions to
achieve the maximum drug load was obtained when pH of the EP solution, contact time between EP and the IL,
and volume of the carrier were 4.6, 11.5 min, and 51.8 µL, respectively. The contour plots are presented in Fig. 5, showing that there was no direct linear relationship between the selected independent variables.
3.4. In-vitro drug release
8
Based on the optimization results 51.8 μL of the IL was placed into a 500 μL EP solution (5 μg mL
-1
) and shaken
for 11.5 min to load the drug. To investigate the drug release, a 20-µL microdroplet loaded by EP was placed into
the channel and the electrotactic drug delivery method was conducted. After the IL microdroplet reached to the
anode, EP was released into the buffer solution inside the hollow fiber. The drug release process was studied by
using buffer solutions with different pH (5.4 and 7.4), at different times. The receptor buffer was withdrawn by a
microsyringe and injected into the HPLC system for measurement of the released EP. Then, the receptor solution
was replenished with fresh phosphate buffer and the new experiment was started. The amounts of released EP
were quantified at 254 nm using a HPLC-UV system, according to a standard calibration graph, obtained by direct
injection of standard EP solutions. The drug release percentage at any specified time was calculated as per Eq. 4.
Drug release (%) =
EP mass in the taken receptor EP mass loaded on the IL microdroplet
The results of the drug release experiments are depicted in Fig. 6. The results demonstrated that 78.2% of the
loaded EP was released from the IL carrier within 10 h at pH 5.4, while it was 54.4% for pH 7.4 at the same
conditions. Better release of EP from the IL at lower pH values is probably due to the protonation of the amine and hydroxyl groups in the EP structure. EP has a basic character (pKa=9.53) and easily protonated at acidic
medium, leads to more positive charge on its structure. Therefore, its solubility in the IL microdroplet with a bulky
cation ([P
6,6,6,14
] ) will be diminished at lower pH values.
4. Conclusions and prospective trends
Epirubicin was loaded with an IL microdroplet and delivered using an in-vitro electrotactic drug delivery system
via a biphasic system. The electrotactic drug delivery experiments were conducted using a microfluidic channel
fabricated using 3D printer. In the presence of an electrochemically generated ion gradient inside the channel,
asymmetric release of [P
6,6,6,14
] from the microdroplet causes its controllable movement. The electrotactic
movement is dependent on applied voltage, drop size, electrolyte concentration, and channel geometry. At the
same conditions, the microdroplets moved more rapidly in half-circular channels than triangular and rectangular
channels. The developed electrotactic drug delivery system showed good capability for delivery of therapeutic
payloads to the sickness sites, leading to an increase of the remedial influence and decrease of side effects. Since
there are different type of IL microdroplets that can be set into motion, smart droplets can be considered as a
promising tool to develop new drug delivery systems. However, much more investigations need to be carried out to find more safe and biocompatible drug carriers to mimic bio-systems or to be used in-vivo.
9
Conflicts of interest
There are no conflicts of interest to declare.
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Figure captions
Fig. 1 Schematic representation of the microfluidic device used for drug delivery studies
Fig. 2 Structure of trihexyltetradecylphosphonium cation (a) and simple planer structure of epirubicin hydrochloride (b)
Fig. 3 Electrical current generated across the microfluidic channel during the electrotactic delivery of EP by applying different voltages to the electrolyte solution (0.01 and 0.001 M)
Fig. 4 Response surfaces plots obtained from the CCD model, a) Vc/time, b) Vc/pH, and c) time/pH
Fig. 5 Contour graphs for the CCD model a) time versus pH, b) Vc versus pH, and C) Vc versus time
Fig. 6 In-vitro release of EP from the IL microdroplet into receptor solutions with different pH, after the electrotactic drug delivery
Recent developments on self-propelled microdroplets, moving controllably in response to an external stimulus
like chemical, electrical, or magnetic field, have opened a new horizon for smart drug delivery investigations. On
the other hand, the new achievements in 3D printing technology has provided a promising option for the
fabrication of microfluidic devices, which is an unrivalled platform for in-vitro drug delivery studies. By
synergizing the features of chemotaxis, 3D printing, and microfluidic techniques a new approach was introduced
to deliver the drug to targeted sites with a well-controlled method and a reasonable speed. A self-propelled ionic
from the microdroplet into an aqueous solution with chloride gradient concentration (created under an external
electrical field) caused the microdroplet to move. The spatial and temporal position of the moving microdroplet
could be controlled by changing the magnitude and polarity of the external electrical field. A piece of hollow-
fiber, fixed next to the anode, was filled with phosphate buffer (as the receptor) and used to remove the drug from
the carrier. The receptor solution was then taken and injected into a HPLC system for quantification of the released
drug. After one-at-a-time optimization of the channel geometry and electrolyte concentration, the experimental
variables affecting the drug loading including contact time, pH, and volume of IMI 28 carrier were optimized via a central composite design (CCD) approach.