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Microfluidic Lung Airway Coculture System with Arrayable Suspended Gels for Studying Epithelial and
Smooth Muscle Cell Interactions
Mouhita Humayun1†, Chung-Wai Chow2.3, Edmond W.K. Young1,4*
1 Department of Mechanical & Industrial Engineering, University of Toronto, Toronto, ON, Canada
† Current location: Department of Biomedical Engineering, University of Wisconsin-Madison, WI, USA
2 Department of Medicine, Faculty of Medicine, Toronto General Hospital Research Institute, University
of Toronto, Toronto, ON, Canada
3 Department of Chemical Engineering, University of Toronto, Toronto, ON, Canada
4 Institute of Biomaterials & Biomedical Engineering, University of Toronto, Toronto, ON, Canada
Keywords: Organ-on-a-chip, respiratory system, suspended microfluidics, air-liquid interface, airway
epithelium, bronchial smooth muscle, goblet cells
* Corresponding author:
Prof. Edmond W.K. Young
Department of Mechanical & Industrial Engineering
Institute of Biomaterials & Biomedical Engineering
University of Toronto, Toronto, ON, Canada
E-mail: [email protected]
Tel.: +1 (416) 978-1521
Fax: +1 (416) 978-7753
2 Abstract
Chronic lung diseases (CLDs) are regulated by complex interactions between airway epithelial cells
(ECs) and airway smooth muscle cells (SMCs), but underlying molecular mechanisms are not well
understood. To advance our understanding of lung pathophysiology and accelerate drug development
processes, new innovative in vitro tissue models are needed that can reconstitute the complex in vivo
microenvironment of human lung tissues. Organ-on-a-chip technologies have recently made significant
strides in recapitulating physiological properties of in vivo lung tissue microenvironments. However,
novel advancements are still needed to enable the study of airway SMC-EC communication with matrix
interactions, and to provide higher throughput capabilities and manufacturability. We have developed a
thermoplastic-based microfluidic lung airway-on-a-chip model that mimics the lung airway tissue
microenvironment, and in particular, the interactions between SMCs, ECs, and supporting extracellular
matrix (ECM). The microdevice is fabricated from acrylic using micromilling and solvent bonding
techniques, and consists of three vertically stacked microfluidic compartments with a bottom media
reservoir for SMC culture, a middle thin hydrogel layer, and an upper microchamber for achieving air-
liquid interface (ALI) culture of the epithelium. A unique aspect of the design lies in the suspended
hydrogel with upper and lower interfaces for EC and SMC culture, respectively. A mixture of Type I
collagen and Matrigel was found to promote EC adhesion and monolayer formation, and SMC adhesion
and alignment. Optimal culturing protocols were established that enabled EC-SMC coculture for more
than 31 days. Epithelial monolayers displayed common morphological markers including ZO-1 tight
junctions and F-actin cell cortices, while SMCs exhibited enhanced cell alignment and expression of α-
SMA. The thermoplastic device construction facilitates mass manufacturing, allows EC-SMC coculture
systems to be arrayed for increased throughput, and can be disassembled to allow extraction of the
suspended gel for downstream analyses. This airway-on-a-chip device has potential to significantly
advance our understanding of SMC-EC-matrix interactions, and their roles in the development of CLDs.
3 Introduction
The human respiratory system is one of the major organ systems of the body, and consists of specific
organs and structures that enable gas exchange between inhaled air from the atmospheric environment
and blood within the circulatory system. The structures of the respiratory system are most often divided
into two regions: (i) the conducting zone, comprised of larger airway tubes that serve the function of
carrying inhaled (or exhaled) air to (or from) the gas exchange region of the lung; and (ii) the
respiratory zone or acinus region, which consists of respiratory bronchioles, alveolar ducts, and alveoli,
and is responsible for actual gas exchange and transport through thin membranous tissue walls that
separate inhaled air from the vasculature.1 Both regions of the lung are susceptible to disease
development, with asthma and chronic obstructive pulmonary disease (COPD) affecting the airways,
leading to airway hyperresponsiveness (AHR), inflammation, and airflow obstruction, and airway tissue
remodeling.2,3 The global burden of asthma and COPD is expected to rise in coming years due to an
aging population and growing exposure to indoor and outdoor air pollution and various allergens.4 Thus,
both fundamental advances in lung biology research and novel therapies for managing lung diseases are
urgently needed.
Development of chronic lung diseases (CLDs) such as asthma and COPD is ultimately regulated
by cell and tissue processes within the airways. Different cell types comprise the airway tissue
microenvironment, including airway epithelial cells (ECs), smooth muscle cells (SMCs), endothelial
cells, and resident immune cells. Of these types, ECs and SMCs have been studied extensively, and have
been shown to play major roles in the pathogenesis of both asthma and COPD.2,3 Importantly, the
communication between ECs and SMCs is a crucial aspect in CLDs triggered by exacerbations such as
air pollutants and pathogens. These stimulants interact directly with the airway epithelium, and lead to
clinical manifestations such as airway narrowing and obstruction that involve SMC proliferation and
4 contraction.2,3 SMC-EC interactions are thus central to elucidating mechanisms in CLDs, and important
for revealing new therapeutic opportunities.
Despite this progress in our understanding of SMC-EC interactions in lung pathophysiology,
many questions regarding the molecular mechanisms regulating lung airway disease processes remain
challenging to address, especially in the context of disease development and exacerbations caused by
pollutant and pathogen exposure. Tackling these challenging questions requires appropriate tissue and
organ models that are convenient, affordable, and more importantly, can accurately represent the
structure and function of the human lung airway microenvironment found in vivo. Most experimental
lung airway models, however, have limitations that do not accurately recapitulate critical aspects of the
in vivo tissue microenvironment. For example, Transwell membrane inserts are used widely for
establishing cocultures between SMCs and epithelia grown on the membrane at the air-liquid interface
(ALI).5,6 However, the membrane is typically made of rigid polymeric materials (e.g., polyester or
polyethylene terephthalate (PET)), which are bio-inert and do not properly mimic the extracellular
matrix (ECM)-laden lamina propria separating the SMCs and ECs within airway tissue. Physiological
airflow may contribute to mechanoregulation of the epithelial layer, but such airflow cannot be readily
applied in typical well-plate or Transwell insert culture formats. Aside from in vitro models, in vivo
animal models are also commonly used in lung airway research studies. Mouse models are especially
valuable due to the wide availability of transgenic models.7 In addition, ex vivo lung slices may be
obtained from mice to allow controlled experimental exposure studies directly on whole tissue
constructs.6 However, while the structure and function of ex vivo tissue slices are more physiologically
relevant than existing culture models, animal tissues inherently differ from human tissues.8 Anatomical,
physiological and immunological differences need to be addressed when translating results from these
tissue sources to human diseases.
5 Recent advances in microengineering and microfluidic systems have enabled researchers to
engineer increasingly more physiological cell culture microenvironments that can provide cells with
various nutrients and physiological stimuli.9,10 In addition, these technologies can be leveraged to build
organ-level tissue structures and geometries with the potential for parallelization and increased
throughput.11,12 Previous microfluidic lung models have been used to investigate the effects of vascular
perfusion, liquid plug flows and cyclic mechanical strain on recapitulating healthy and damaged lung
epithelium.13,14 Recently, much more attention has focused on studying the synergistic effects of
multiple lung airway tissues in the progression of lung inflammatory disorders, particularly the lung
epithelium and the underlying vasculature, on inflammatory responses to pathogens, neutrophil
recruitment and molecular secretion profiles, providing insights into the mechanisms of inflammatory
lung disorders.14,15 In these models, the two tissue types are compartmentalized in vertically stacked
chambers separated by a thin porous polyester membrane coated with exogenous ECM to allow
crosstalk between the tissue types. These models demonstrate the potential for organ-on-chips, and have
opened new opportunities to extend microfluidic technologies to achieve even more complex and
physiologically relevant models.11,15,16 Despite their transformative potential, these designs still have
room for significant advancement in biological complexity, ease of fabrication, increase in throughput,
and additional sample manipulation. The biologically inert polymeric membranes used in current
microchip designs have demonstrated support of cell attachment and growth, but are not representative
of the structure and stiffness of native ECM. Furthermore, coculture between epithelial and
microvascular endothelial cells have been shown, but coculture between ECs and SMCs have yet to be
demonstrated. This is crucial for studies of airway injury and repair after pollutant exposure because
SMCs rather than endothelial cells are present in large and medium airways; endothelial cells are only
present in the smallest airways of the respiratory tree.1 Additionally, current designs are not often as
scalable as desired, and can be time consuming and laborious to fabricate due to challenges of
6 incorporating the polymeric membrane. Therefore, to push the boundaries of organ-on-a-chip
technology, new innovative designs are necessary that can demonstrate coculture of other cell types,
enable transition to biocompatible growth surfaces, and offer increased throughput and reduced
fabrication time to maximize experimental efficiency.
Here, we describe a novel microfluidic cell-based model that offers three unique advantages
from other recently developed microfluidic systems. First, this model uniquely recapitulates airway EC-
SMC interactions in microfluidic cell culture, which can be distinguished from the other models that
have focused on epithelial-endothelial interactions in the small airway. Achieving appropriate SMC
morphology and function on the chip in co-culture with airway epithelia is important for studying
airway constriction and matrix remodeling. Second, rather than separating top and bottom microfluidic
compartments with a biologically inert polyester membrane, our design employs a unique suspended
hydrogel to separate top and bottom compartments. The advantage of a suspended hydrogel is its
biocompatibility, and the fact it can be comprised solely of naturally derived matrix components. Third,
we demonstrate the use of a thermoplastic construction with an arrayable design, features which enable
potential increases in throughput and allow a path towards manufacturability.
7 Materials & Methods
PMMA Micromilling
Our lung airway microdevice (or “airway-chip”) was fabricated from two 1/16″ thick and one 3/32″
thick flat pieces of poly(methylmethacrylate) (PMMA) (#8560K173 and #8560K183, McMaster-Carr,
Elmhurst, IL, USA), which were micromilled with features using in-house protocols.17,18 The device
features were modeled in SolidWorks (Dassault Systèmes, Velizy-Villacoublay, France) and converted
to G-code in SprutCAM (SprutCAM, Naberezhnye Chelny, Russia) for computer numerical control
(CNC) milling. Microfeatures were generated using a Tormach PCNC 770 vertical milling machine
(Tormach, Waunakee, WI) with 4-flute carbide endmills of the following sizes: 1/16″ (1.5875 mm,
#01982156), 3/64″ (1.1906 mm, #07765431) and 0.015″ (381 µm, #37289501) (MSC Industrial Supply
Co., Melville, NY, USA). Key microchannel feature dimensions were as follows: (1) for the top layer of
the device, the channel designed for air exposure (for ALI) was 1.2-mm deep, 5-mm wide by 10-mm
long; (2) for the middle layer of the device, the main channel was 0.7-mm deep, 5-mm wide by 10-mm
long, and the inner microchannel for the suspended hydrogel was 0.65-mm deep, 2.0-mm wide by 6-mm
long. The exposed region to the media reservoir was 0.85-mm wide and 4-mm long; (3) for the bottom
layer, the lower media reservoir channel was 1.2-mm deep, 1.6-mm wide x 12.5 mm long. The inlet and
outlet ports in each layer were 2.0-mm and 2.75-mm in diameter, respectively.
Solvent-Assisted Thermal Bonding
To bond the layers of PMMA, we employed a solvent-assisted thermal bonding method developed in
our lab.19 Briefly, 99% ethanol (Commercial Alcohols Inc., Toronto, ON, Canada) was pipetted between
two layers of PMMA, aligned and pressed with 1000 lbf at 70 °C for 1 min with a Carver Automatic
Hydraulic Laboratory Press (#3889, Carver Inc., Wabash, Indiana, USA). The middle and bottom
8 PMMA layers were bonded first, followed by the bonding of the top layer to the middle layer. See Wan
et al. for step-by-step procedures.19
Standard Cell Culture
Human airway epithelial cells (Calu-3) were obtained from ATCC (HTB-55, Manassas, VA, USA) and
maintained in Minimum Essential Medium α (MEM-α) (#12561-056, Thermo Fisher Scientific,
Burlington, ON, Canada) supplemented with 10% fetal bovine serum (FBS) and 1%
penicillin/streptomycin (P/S) (Thermo Fisher Scientific). Human bronchial smooth muscle cells
(hBSMCs) were obtained from Lonza (#CC-2576, Cedarlane Labs, Burlington, ON, Canada), and
maintained in smooth muscle basal medium (SmBM) supplemented with BulletKit (SmGM-2, Lonza)
and 1% P/S. All cells were maintained in an incubator at 37 °C and 5% CO2. Calu-3 cells were used
between passages 11 to 14, and hBSMCs were used between passages 9 to 12.
Hydrogel Preparation
Rat-tail Type I collagen (CACB354249, Corning) and growth-factor-reduced Matrigel (CACB354320,
Corning) were purchased from VWR (Mississauga, ON, Canada). Collagen was mixed with 20% 0.5-N
NaOH (Sigma-Aldrich, Oakville, ON, Canada) to neutralize the solution, and then incubated at 4 °C for
1 h to allow the formation of thicker collagen fiber structures.20 This collagen solution was then mixed
with Matrigel for preparations of the mixed hydrogels that were tested.
Loading and Culture in Microdevice
To disinfect the lung microdevice, all the internal chambers of the device were flushed at least three
times each with 70% ethanol and 1X PBS. During the last wash, a thin film of PBS was left to render the
walls of the hydrogel suspension channel hydrophilic. To improve adhesion of the hydrogel to the
9 channel walls, 100 µg/mL bovine fibronectin (FN) (F1141, Sigma Aldrich) was pipetted into the
channel immediately after the last PBS flush, and the device was incubated at room temperature for 30
min. After incubation, the FN was aspirated out of the hydrogel suspension channel and a thin coat
along the channel walls was left to improve capillary flow of the hydrogel. The hydrogel was dispensed
into the hydrogel suspension channel, and polymerized in a humidified incubator at 37 °C, 5% CO2 for 1
h, resulting in a suspended polymerized exogenous hydrogel (Fig. 1E). After polymerization, the
chambers of the devices were filled with Calu-3 cell culture media and incubated overnight to keep the
hydrogel hydrated.
Airway Epithelial Cell Culture
Calu-3 cells were seeded as an 80-µL suspension at 3000 cells/µL on top of the suspended hydrogel of
the lung microdevice. The cells were incubated at 37 °C and 5% CO2 overnight in the device, and the
media was replaced the next day to remove unattached cells from the top compartment. The top and
bottom chambers were filled with Calu-3 cell culture media during normal epithelial cell culture and
replenished every 24 h. Cells were cultured for 14 days prior to immunostaining. For air-liquid interface
(ALI) cultures, Calu-3 cells were cultured on the suspended hydrogel for 5 to 7 days before removing
the media from the top chamber and exposing the cells to air. During ALI cultures, the apical surface
was rinsed with PBS daily and maintained at ALI for 3 weeks prior to immunostaining. Calu-3 media
was used in the basolateral side (bottom) compartment and was replenished every 24 h.
Coculture of Lung Epithelial and Smooth Muscle Cells
Calu-3 cells were cultured on the suspended hydrogel, as detailed above, for 7 days prior to seeding of
hBSMCs. At Day 7, hBSMCs were seeded to the bottom chamber as a 50-µL suspension at 750 cells/µL
and the device was turned over to facilitate adhesion of hBSMCs to the underside of the suspended
10 hydrogel. The device rested on two PDMS posts during upside-down culture, and was incubated for 4 h.
Subsequently, the media was replaced in both compartments to remove any unattached cells. The co-
culture was maintained for 1 week prior to immunostaining, and media in both the top and bottom
compartments was replaced every 24 h.
Device Disassembly and Sample Removal
Cells inside the lung microdevice were fixed with 4% paraformaldehyde (Commercial Alcohols Inc.,
Toronto, ON, Canada) for 20 min and permeabilized with 0.1% Triton X-100 (Sigma Aldrich) for 10
min. Cells were washed with PBS three times between each step by flushing the solution through the top
and bottom chambers of the device. To extract the suspended hydrogel sample with cultured airway cells
intact, the top and middle layers of the PMMA device were separated using a single-edge steel razor
blade (Fig. 2A). The blade edge was inserted between each layer at the corner of the device, and gently
moved along the four device edges to slowly open the PMMA-PMMA bonds between layers. Once the
top layer was removed, a scalpel was used to separate the hydrogel from the inner walls of the
microchannel. The hydrogels were then gently rinsed with PBS using a pipette until the hydrogels
completely detached from the device. The released hydrogel was subsequently suspended in PBS in a
0.6-mL centrifuge microtube.
Immunostaining
Standard immunofluorescence staining procedures were performed on hydrogel samples while they
were suspended in 0.6-mL centrifuge microtubes. To visualize tight junctions, samples were blocked
with 3% bovine serum albumin (BSA) & 0.1% Tween-20 in PBS (Sigma-Aldrich), and stored at 4 °C
overnight. ZO-1 primary antibody (Rabbit anti-ZO-1 pAb; Life Technologies, Carlsbad, CA, USA) was
diluted to 10 µg/mL in the above-mentioned blocking buffer and added to the sample. The samples were
11 stored at 4 °C for 48 h. Samples were then labeled with AlexaFluor 568 goat anti-rabbit IgG (10 µg/mL
dilution in blocking buffer; Life Technologies, Carlsbad, CA, USA) and Hoechst 33342 (1 µg/mL
dilution in blocking buffer) at 37 °C, 5% CO2 for 2 h. Samples were flushed with PBS prior to and after
labeling with secondary antibody before imaging. The same procedure was used for α-smooth muscle
actin (α-SMA) staining in hBSMCs, where mouse anti-α-SMA mAb (Sigma, St. Louis, Missouri, USA)
was diluted at 10 µg/mL in blocking buffer and added to the sample. AlexaFluor 488 goat anti-mouse
IgG (Life Technologies, Carlsbad, CA, USA) at 10-µg/mL dilution in blocking buffer was used for
labeling. MUC5AC primary antibody (Mouse anti- MUC5AC mAb; Life Technologies, Carlsbad, CA,
USA) at 20-µg/mL dilution was used to label mucin-producing airway epithelial cells, and conjugated
with AlexaFluor 488 goat anti-mouse IgG using the above-mentioned procedure. To identify F-actin in
both airway cells, AlexaFluor 488 phalloidin (Life Technologies, Carlsbad, CA, USA) at 2.5% (v/v) in
blocking buffer was used.
Imaging and Data Analysis
For imaging, excess fluorescent labeling solution was removed with PBS and samples were mounted on
a 1.2-mm thick 75 x 25 mm glass microscope slide (Fig. 2C). Images were obtained using a fluorescence
microscope (EVOS FL Auto Cell Imaging System). For hBSMC alignment analysis, Hoechst (nuclear)
stained images were analyzed using the ImageJ Orientation J plugin. To measure the angles of each
nuclei, an ROI was selected covering >90% of each stained image.
Statistical Analysis All values are presented as mean ± standard deviation. In total, four independent experiments (n = 4)
were conducted, which means that results from four separate culture systems were used for each
condition on each day for each of the two different cell types. Single-variable analysis of variance
12 (ANOVA) was used for multiple comparisons within a study (e.g., effect of hydrogel on cell adhesion of
each cell type), and multiple comparison tests were performed using Tukey’s method. To analyze the
effects of hydrogel compositions and culture time on the area covered by adhered cells, two-way
ANOVA was performed followed by Sidak’s multiple-comparison test. Statistical differences were
considered significant for p < 0.05.
13 Results and Discussion
Lung Microdevice Design and Fabrication
To recapitulate the in vivo microenvironment of the lung airway within the conducting zone, we
considered the functional roles of major cellular constituents of the airway tissue. The tissue structure of
the airway includes a pseudostratified differentiated epithelium lined with a band of smooth muscle cells
(SMCs) separated by a thin layer of connective tissue called the lamina propria (Fig. 1A).21 These tissue
layers together perform critical roles in the adaptive and innate immune system responses of the airway
that drive development of CLDs.22
Given the roles of these tissue layers in airway cellular responses, we incorporated in our model
a bronchial epithelial cell type (Calu-3 cell line), and a primary human bronchial smooth muscle cell
(hBSMC) type, cultured on opposing sides of a suspended hydrogel. Our device design consisted of
three vertically stacked compartments, fabricated from three milled PMMA layers, where the top and
bottom chambers compartmentalized two different airway cell types and were separated by a suspended
hydrogel in the middle compartment (Fig. 1B-D). Separate inlet and outlet ports were incorporated to
accommodate daily media replenishment for the cells in both top and bottom compartments (Fig. 1D).
The inlet port for the epithelial cells cultured in the top chamber were located above the hydrogel port to
allow easy access to the hydrogel chamber during loading. The outlet port of the top chamber was
located away from the hydrogel outlet port to avoid disturbing the hydrogel during media changes.
14
Figure 1. Lung airway-on-a-chip device design and fabrication. (A) Schematic of airway lumen tissue structure including a thin ECM layer (lamina propria) sandwiched by a ciliated airway epithelium on one side and aligned bronchial smooth muscle cells on the other. (B) Photograph of an assembled lung airway microdevice with the top, middle and bottom compartments highlighted in yellow, green, and red, respectively (scale bar = 10 mm). (C) Lung airway microdevice schematic design showing three vertically stacked layers of PMMA with a top airflow chamber, middle suspended hydrogel chamber and bottom media reservoir. (D) Exploded view of the three PMMA layers that comprise the microdevice, and assembled view of the solvent bonded PMMA layers. (E) The suspended hydrogel is introduced into the middle layer by simple pipette loading. Magnified image of device cross-section showing hydrogel loaded with red-fluorescent microparticles, supported by surface tension between hydrogel and the surfaces of the two protruding ledges (scale bar = 500 µm).
15 PMMA was chosen as our device material because of several desirable properties, including
bioinertness, compatibility with existing mass production infrastructure, low material cost, and optical
transparency. While PDMS has been a popular choice for biomicrofluidic systems, we (and others) have
observed how PDMS can absorb small hydrophobic molecules from solution, which can influence cell
biology readouts, and lead to unintended biases.23,24 Thus, for certain biomicrofluidic applications,
thermoplastics may be more preferable than PDMS.25 Our combined approach of micromilling and
solvent bonding allowed us to align various plastic layers during the bonding process repeatedly. With
micromilling, microscale features can be machined into plastic substrates within the timeframe of
minutes to hours, depending on the complexity of the design. Moreover, we have previously
demonstrated that these milled plastic layers can be solvent-bonded with high bond strength by
employing retention grooves, thereby achieving uniform bond coverage between layers and preserving
the microscale features that have been milled into the layers.19,23,26 This process helped to ensure a leak-
free bond in our devices, which enabled us to incorporate an array of culture systems on a single chip
without cross-contamination.
To incorporate a more biologically relevant component in our system, we leveraged the concepts
of suspended microfluidics27 to integrate a suspended hydrogel rather than a polymeric membrane to
separate the two cell types, and thus attempt to model the lamina propria using matrix proteins in a
microfluidic system. The thickness of the suspended hydrogel can be modified by changing the
thickness of the protruding ledges of the middle plastic layer. To our knowledge, this is the first
demonstration of airway epithelial and smooth muscle cell co-culture on opposing sides of a suspended
hydrogel in a microdevice. By simply pipetting a thin layer of hydrogel between the two protruding
PMMA ledges, the pre-polymerized hydrogel flowed by capillary action along the ledges, thus creating
a thin film of gel with two open interfaces for cell culture (Fig. 1E).
16
Arrayability and Sample Extraction
Using the thermoplastic fabrication techniques described above, we fabricated an array of 12
independent culture systems on a single PMMA device, which matched the format of a 75 x 25 mm
microscope slide (Fig. 2; same channel design with increased array throughput compared to Fig. 1B).
The array increased throughput without risking cross-contamination between adjacent systems due to
leak-free solvent bonding. Cells in culture can be monitored for viability inside each system at various
time points via immunofluorescence staining within the optically clear device. To demonstrate the
versatility of our device in enabling downstream sample processing, we disassembled our device and
extracted cell culture samples with minimal disruption to the sample itself by lodging a single edge razor
blade and severing the interfacial bonds between the top and middle PMMA layers (Fig. 2A-B). The
samples consisted of both cell types adhered to the hydrogel material in between, and effectively
represented a small engineered “tissue” sample that can be readily handled and manipulated (Fig. 2C).
For endpoint analysis, we fixed and permeabilized the cell cultures on the suspended hydrogel inside the
device prior to disassembling the chip, ensuring that the cells were not affected during device
disassembly. After detaching the top and middle layers of the device, the suspended gels in the middle
layer were easily detached from the device walls with tweezers, and then submerged in PBS for
immunofluorescence staining (IFS) (Fig. 2C-D). With this approach, extracted culture samples from
each system can be combined or individually stained in an Eppendorf tube by flushing the samples with
IFS solutions. Therefore, our device fabrication approach enabled simple extraction of culture samples
with minimal damage to the samples, and can allow further downstream sample processing for DNA,
mRNA, protein and molecular analyses if desired to assess more complex cellular functions.
17
Figure 2. Sample gel extraction for immunofluorescence staining. (A) Photograph of a lung airway microdevice with 12 arrayed systems after sample fixation. The top and middle layers of the device were detached by lodging a single edge blade between the layers. (B) Photograph of a disassembled device after sample fixation. (C) A fixed sample (airway cells and ECM hydrogels) that was detached from the side walls of the channels from the middle PMMA layer using a sharp tip of a scalpel. (D) Stained samples on a microscope slide, ready for microscopy.
18
Cell Adhesion to Hydrogel
We studied different hydrogel mixtures to determine an optimal gel composition that could
support co-culture of both Calu-3 cells and hBSMCs over 7 days. Specifically, we chose to test mixtures
of Type I collagen (Col-I) and Matrigel because of the abundance of Col-I in ECM, and because
Matrigel is known to contain basement membrane proteins and is widely used in in vitro assays due to
its ability to promote adhesion.28,29 We tested the adhesion of both airway cell types on four different
hydrogel compositions: (i) Col-I alone; (ii) Matrigel alone; (iii) a mixture of 6 µg/µL Col-I and 3 µg/µL
Matrigel (i.e., referred to below as “high Col-I mix”); and (iv) a mixture of 6 µg/µL Matrigel and 3
µg/µL Col-I (i.e., referred to below as “high Matrigel mix”). Calu-3 cells and hBSMCs were cultured
and stained for nuclei after 2 days and 7 days, and either counted for hBMSCs, or measured for area
coverage in the case of Calu-3s.
After 2 days, Calu-3 cells formed networks with neighboring cells and established partial
monolayers on Col-I, the high Col-I mix and the high Matrigel mix, covering large regions of the
hydrogel surface area. Cells displayed the least area coverage on Matrigel alone (Fig. 3A-B).
Morphological analysis of Calu-3 cells after 2 days confirmed significant statistical differences (**** p
< 0.0001) in area coverage by cells on Matrigel versus all other hydrogel solutions. Interestingly, after 7
days, cell detachment was observed on the Col-I and the high Col-I mix hydrogels, but became more
interconnected as a monolayer on the high Matrigel mix. The low adhesion behavior on Matrigel alone
remained unchanged over the 7-day culture period. After 7 days, significant differences were observed
in area coverage between the high Matrigel mix and all other compositions tested in this study (**** p <
0.0001). No significant difference (p > 0.05) was found in area coverage between Col-I and the high
Col-I mix after 7 days. Area coverage on the high Matrigel mix remained above 80% from Day 2
through to Day 7 (Figure 3B).
19
Figure 3 Calu-3 epithelial cell and hBSMC adhesion on hydrogel in a PMMA lung airway microdevice. Adhered cells were stained with Hoechst 33342 to label the nuclei and observed under fluorescence microscopy on Days 2 and 7. (A) Fluorescent images represent Calu-3 cell adhesion on Type I Collagen alone (Col-I); mixture of 6µg/µL Col-I and 3 µg/µL Matrigel (C6:M3); mixture of 6 µg/µL Matrigel and 3µg/µL Col-I (M6:C3); and Matrigel alone (Matrigel). Dashed lines represent edge of hydrogel. Scale bar = 400 µm. (B) Comparison of area covered by adhered Calu-3 cells (%) on different hydrogel compositions after 2 d and 7 d. Data presented as mean ± SD (n = 4, **** p < 0.0001). (C) Cell adhesion and growth of hBSMCs cultured for 7 d on different hydrogels. Graphs represent a comparison of the number of attached hBSMCs per mm2 of hydrogel surface inside a PMMA microfluidic device after 2 d and 7 d of culture. Data presented as mean ± SD (n = 4).
20 For primary hBSMCs, cells appeared to be uniformly dispersed (without networks) on Col-I, the
high Col-I mix and the high Matrigel mix. Similar to Calu-3s, hBSMCs appeared to form aggregates in
small regions on Matrigel alone instead of evenly across the surface. Combined with the results obtained
for Calu-3 adhesion, the low adhesion of both cell types on Matrigel alone warranted excluding Matrigel
alone from further study in our device. After 2 days on the other three gel compositions, hBSMCs
appeared to be well dispersed on all three hydrogels. Quantitative analysis of attached cells per mm2
revealed no significant differences (p > 0.05) between the three hydrogel compositions after both 2 and
7 days (Fig. 3C). However, cells appeared to be more confluent after 7 days, suggesting some
proliferation may have occurred. Nuclei count confirmed these observations of proliferation from Day 2
to Day 7, as well as no significant differences in cells attached per mm2 between the three hydrogels
tested on both Day 2 and Day 7.
While previous studies have shown increased Calu-3 cell adhesion, proliferation and confluent
monolayer formation on Col-I-coated substrates compared to non-coated substrates,30,31 our results
suggested that cell adhesion may be quite different when cultured on a Col-I hydrogel of ~500-µm
thickness compared to a more conventional thin coating of Col-I. Furthermore, this indicated that aside
from integrin signaling between cells and the ECM that are known to influence cell adhesion, other
factors such as substrate stiffness, surface topography and availability of growth factors may also play
important roles. The potential impact of Col-I gel stiffness on Calu-3 epithelial adhesion may thus
require further investigation.
In addition, airway smooth muscle cell adhesion studies on 2D substrates in the past have
suggested that interstitial ECM proteins like Col-I promote proliferation and expression of less
contractile phenotypes while basement membrane proteins stimulate expression of a differentiated
contractile phenotype and inhibits proliferation.32,33 This is due to the effects of laminin, a constituent of
Matrigel, and its primary integrins that inhibit airway SMC growth.34 Consistent with these findings, we
21 found that hBSMCs seeded on Matrigel did not adhere and spread over a 7-day culture period compared
to other hydrogels tested. Interestingly, hBSMC adhesion was not affected by the hydrogel mix with
high Matrigel concentration. A possible explanation for this observation is that the primary integrin
receptors, more specifically α1β1 and α2β2, that mediate growth inhibition of laminin also interact with
Col-I 35, which is present in our hydrogel mixture. These interactions may have prevented the growth
inhibitory effects of laminin on hBSMCs, leading to improved cell adhesion compared to Matrigel
alone. This may also explain previous results that showed successful endothelial cell adhesion on a
mixture of Col-I and Matrigel within a microfluidics context.36
Together these results showed that a mix with Col-I and high Matrigel concentration facilitated
increased initial and sustained Calu-3 cell and hBSMC adhesion over a 7-day culture period.
Specifically, our results showed that Col-I was important for initial cell adhesion to matrix, whereas the
ECM proteins in Matrigel facilitated sustained cell adhesion over a longer culture period. Based on these
findings, the high Matrigel mix was used for all subsequent work in this study.
ALI Culture Protocol and Characterization
Calu-3 cells were cultured on the top surface of the suspended hydrogel for 14 days, with daily media
replacements on both apical and basolateral sides, before exposing to ALI for more than 2 weeks (17
days) (Fig. 4A). During ALI culture, only the basolateral media was replaced, and the apical side was
flushed with 1X PBS every 3 days to remove accumulating mucus. To examine both intracellular and
barrier structures of the epithelium, we immunostained for F-actin and ZO-1 tight junctions after 7 and
14 days, respectively (Fig. 4B). F-actin staining revealed strong localization around the periphery of
each epithelial cell, consistent with a typical cell cortex found with cells within a monolayer (Fig. 4B).
ZO-1 staining showed cobblestone morphology and well-defined polygonal rings localized to the outer
periphery of the cells. This morphology was less pronounced after 7 days when ZO-1 was more diffuse,
22 and more pronounced after 10 days of submerged culture when both F-actin and ZO-1 rings were clearly
visible (Fig. 4B, 14 days shown). This type of morphology is common in Calu-3 monolayers, and is
indicative of stable barrier properties formed in these monolayers.37,38 We chose to expose epithelial
cells to ALI after 14 days of submerged culture to ensure stable tight junction barrier formation prior to
ALI exposure.
At 31 days (i.e., 14 days of submerged culture followed by 17 days of additional ALI culture),
Calu-3 cultures were fixed, extracted and immunostained for MUC5AC, a goblet cell marker and a
protein abundant in the mucosal layer of the airway. We observed that dispersed groups of Calu-3 cells
on the monolayer stained positive for MUC5AC expression (Fig. 4C), suggesting that our system was
capable of accommodating ALI culture conditions that can facilitate airway epithelial differentiation into
goblet cells. This is consistent with cellular compositions typically found in native small airways.
Furthermore, to ensure that the tight junctions that appeared by 14 days were not compromised during
ALI culture, we co-stained the culture for ZO-1, and found similar tight junction formation after 31 days
(Fig. 4C). These results showed that the airway epithelium in our lung device can establish tight
junctions between cells, and differentiate into mucus-producing goblet cells. We also demonstrated the
ability of our device to accommodate long-term airway epithelial cell cultures, including ALI culture
that can last several weeks (> 4 weeks) without losing barrier structure or differentiation capacity.
23
Figure 4. ALI culture protocol and characterization of Calu-3 cell cultures on suspended hydrogel. (A) Protocol for air-liquid interface (ALI) culture of epithelial cells (dotted line represents longitudinal cross-sectional plane in illustrations). Cells are seeded and cultured under submerged culture conditions for 14 days prior to ALI exposure for an additional 17 days with daily media replacement. (B) Immunofluorescence image of monolayers formed by cultured Calu-3 cells after 7 and 14 days. Green: F-actin; Red: ZO-1 tight junctions; Blue: Hoechst nuclei (scalebar = 20 µm). (C) Immunofluorescence image of Calu-3 cells in monolayer expressing goblet cell marker MUC5AC (green) and ZO-1 tight junctions (red) (scalebar = 50 µm).
24 Coculture Protocol and Characterization
For co-culture of both airway ECs and SMCs, Calu-3 epithelial cells were first seeded and cultured on
the suspended hydrogel for 7 days under submerged culture conditions. Subsequently, hBSMCs were
seeded in the bottom compartment, and the device was flipped upside-down for 4 h to allow the
hBMSCs to settle due to gravity and adhere to the underside of the gel. After adhesion, the device was
flipped back to the upright position for the remainder of the culture period, i.e., an additional 14 days
prior to fixation, extraction and staining (Fig. 5A). We performed co-culture characterization to
highlight phenotypic differences between the two cell types and to determine if the characteristic
properties of Calu-3 monolayer and hBSMCs can be reproduced and maintained under coculture
conditions. At 14 days, the coculture samples were co-stained for ZO-1, nuclei and alpha smooth muscle
actin (α-SMA), a phenotypic marker for SMCs. Similar to the monocultures, Calu-3 cells displayed
well-defined polygonal rings of ZO-1 localized to the periphery of the cells under co-culture conditions,
indicating that hBSMCs did not interfere with epithelial tight junction formation (Fig. 5B). This
morphology was observed uniformly across the Calu-3 monolayers in coculture. Additionally, hBSMCs
were positively stained for α-SMA (Fig. 5C). Confocal 3D images of our cocultures stained with both α
-SMA and ZO-1 displayed a clear distinction between the two cell types separated by the suspended
hydrogel (Fig. 5D-E). Importantly, the integrity of the suspended hydrogel appeared undisturbed
throughout the culture, extraction and staining processes.
25
Figure 5. Protocol for coculturing Calu-3 epithelial cells with hBSMCs (dotted line represents cross-sectional plane in illustrations). (A) Protocol for coculture Calu-3 cells and hBSMCs (dotted line represents cross-sectional plane in illustrations). Calu-3 cells were seeded and cultured under submerged culture conditions for 7 days prior to hBSMC seeding. The device was flipped over for hBSMC adhesion (4 hours) and cocultured with Calu-3 cells in the upright positions for 7 more days with daily media replenishment. (B) Immunofluorescence image of monolayers formed by cultured Calu-3 cells after 14 days of total culture and 7 days coculture (Green: F-actin; Red: ZO-1 tight junctions), scalebar = 20 mm). (C) Immunofluorescence image of alpha smooth muscle actin (α-SMA) in hBSMCs (Red: α-SMA; Blue: Hoechst nuclei) cocultured with Calu-3 cells for 7 days), scalebar = 50 mm. (D) 3-D rendered confocal laser micrograph of Calu-3 and hBSMC coculture on two sides of the suspended ECM hydrogel. (E) Confocal laser micrograph of coculture sample viewed from the x-z plane. (F) Polar graphs showing frequency distribution of the angle of hBSMC alignment with respect to the inlet-outlet axis of the microfluidic channel. Radial (horizontal) axis represents the frequency of angle of alignment. Each bar represents a 10°-range. Arrows point to the outlet of the microfluidic channel. Polar graphs represent frequency distribution for one representative sample on Day 4 (top panel) and Day 7 (bottom panel) of co-culture. A narrower distribution is observed for 7-day co-culture compared to a 4-day co-culture (n = 134). Mean alignment angles and standard deviations of each system are indicated on the top right corner of each graph for each system.
26 Interestingly, in our analysis of α-SMA in hBSMCs, the actin stress fibers appeared to be
predominately oriented in the direction parallel the length of the channel (Fig. 5F). Quantitative
assessment of angle of orientation revealed that this preferential alignment was more significant with
longer culture times and higher cell densities. Actin fiber orientation angles were measured and the
frequency distributions were plotted in polar graphs (Fig. 5F). A total of 134 cells were analyzed for
their fiber orientations from each image obtained and grouped into 10º bins. After 4 days of co-culture,
cells appeared to be more randomly oriented, after 7 days of co-culture, however, a narrower distribution
of alignment angles were observed (Fig. 5F). The angle of orientation spanned from the northwest
direction to the northeast direction with a mean alignment angle of 78º and standard deviation of 43º for
our representative sample of 4-day co-culture. For a 7-day culture sample, a mean alignment angle of
73º and standard deviation of 21º was observed. These observations were consistent throughout replicate
experiments for each day.
An important distinction between the expansion of Calu-3 epithelial cultures in the monoculture
setup (Fig. 4) and coculture setup (Fig. 5) is that there is crosstalk between two different cell types that
may influence the growth and development of these cultures. Here we show that hBSMC coculture with
Calu-3 cells in our lung device did not affect the ability of the epithelial cells to from tight junctions.
Furthermore, hBSMCs appeared to align along the longitudinal direction of the device with extended
culture time, taking on a morphology similar to that seen in vivo.
With this device design, coculturing of SMCs and ECs, hydrogel extraction, and costaining for
various phenotypic markers can easily be done with our chip without cross-contamination between
culture systems. Importantly, the ability to remove the cell-laden hydrogel from the device – with both
cell layers intact and without disruption of the gel – enables possible downstream biological analyses to
study gene and protein expression of either cell type, as well as matrix deposition and remodeling events
that are critical to understanding airway disease development. Of particular interest is the future
27 potential to incorporate airflow over the airway epithelium at the ALI of the upper chamber, which is
possible due to the geometry of the microchannel features with the current design.
Conclusion
We have developed a plastic airway-on-a-chip device for culturing airway ECs and SMCs on opposing
sides of a suspended hydrogel, and performed preliminary experiments to demonstrate long-term
viability and assess its fitness as an in vitro lung airway model. We used a combined micromilling and
solvent bonding approach to fabricate a thermoplastic-based platform that is suitable for making arrays
of culture systems in a single microfluidic device within a matter of hours. The key feature of the design
is the use of “open microfluidics” to suspend a thin layer of hydrogel that can support long-term culture
of both airway ECs and SMCs. The device was also designed to accommodate ALI culture, which was
shown to lead to the differentiation of airway epithelial cells into mucus-producing goblet cells inside
the microfluidic device. As part of preliminary assessment, the device was used to determine an optimal
hydrogel composition that promoted airway cell adhesion and proliferation. In monoculture, ECs
displayed cobblestone morphology with tight junction proteins, and stained positive for the goblet cell
marker MUC5AC under ALI culture conditions without the loss of tight junction markers. In coculture,
ECs retained tight junctions in monolayer, while SMCs appeared highly aligned and expressed α-SMA
proteins, which are critical for motility and contractility functions. These findings demonstrated the
usefulness of the designed platform for its ability to accommodate long-term SMC-EC coculture, to
allow immunofluorescence staining and sample handling without disrupting the matrix or cells, and to
offer increased throughput due to the arrayability of the design. The platform has significant potential
for use as a lung airway tissue model that can incorporate other microenvironmental cues, including
airflow and pollutant exposures for studying mechanisms associated with CLDs.
28 Conflict of Interest
There are no conflicts of interest to declare.
Acknowledgements
We acknowledge financial support from the Natural Sciences and Engineering Research Council of
Canada (NSERC) Discovery Grant, and from the Canadian Lung Association / Ontario Thoracic Society
Grant-in-Aid Award to EY.
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