All the current lung-on-a-chip devices are aiming to recapitulate the complexity of the lungs to some extent to offer an advanced cell-culture model for drug discovery purposes. They usually consist of three components: (1) an air channel where epithelial cells can be grown and transferred later to air-liquid interface, (2) a “blood” channel that serves as the vascular part of the chip, and (3) a semi-permeable membrane to separate the air and fluidic channels. Together,  all these new features can offer an in vitro model that has these advantages over more conventional in vitro models: (1)emulating the complexity of the lungs to some extent for instance by the integration of a stretchable membrane, (2)adding the respiratory breathing motions to the model either by stretching or the deflection of a thin membrane, (3) having “blood” and airflow in the apical and basal sides which can induce shear stress on cells and present a more realistic model than the other static in vitro models, and (4) the feasibility of using an extracellular matrix (ECM) based material as the membrane or added to the polymeric membrane. In some models, two sides of channels have been included to stimulate breathing by generating cyclic pneumatic pressure. This is helpful when the focus of the model is alveolar. Otherwise, a cyclic flow in the lower channel would be sufficient to mimic the physiological shear stress on cells. In addition, some biochemical factors such as proteins (collagen and fibronectin)or growth factors can be added to the chip (coated on both sides of the membrane) to recapitulate the cellular microenvironment.Polydimethylsiloxane(PDMS) is the most common material in the fabrication of these lung-on-a-chip devices since PDMS is optically transparent, biocompatible, easy to use, and flexible. Lung-on-a-chip devices aim to provide a tool for studying a drug. Besides, these devices can be used to simulate a lung disease such as inflammation, asthma, lung cancer, pulmonary fibrosis, or lung injuries and simulate the lung cells/tissue interactions at a “realistic” scale by controlling their microenvironment.

To begin with, we should mention the device designed and developed by Huh et al. in which mechanical strain was applied to stretch a porous PDMS membrane. They showed that the mechanical strain facilitated the uptake and transfer of nanoparticles by epithelial and endothelial. These results were confirmed by observing the similar effects in the whole mouse lung. The same research group later successfully modeled a human disease model-on-a-chip for pulmonary edema. This group led by DE Ingnercan be considered as one of the most dynamic teams in the field that has introduced various inventions not only to lung-on-a-chip devices but also to other organ-on-a-chip devices. In another study, they developed a small airway-on-a-chip and lined the airway with epithelial cells from patients suffering from chronic obstructive pulmonary disease. Using viral and bacterial infections, they could model the disease in their device. Moreover, they fabricated a microfluidic chip containing embedded electrodes for measuring trans-epithelial electrical resistance (TEER), which is used to monitor and quantify the integrity of cultured epithelial under statistic conditions. Later, they developed a lung alveolus-on-a-chip to model pulmonary thrombosis by co-culturing epithelium and endothelium. In another interesting work, KL Sellgren et al. designed a chip for culturing primary endothelial, epithelial, and lung fibroblasts to model the human airway. Polytetrafluoroethylene (PTFE) and polyester (PET) membranes were integrated between different layers of the chip. This work showed the feasibility of culturing primary airway epithelial cells with lung fibroblast and endothelium while providing appropriate compartmentalization in the chip. Such a triple co-culture with the capability of perfusion cannot be provided by the current commercial in vitro models. Other researchers also tried to integrate electrodes in a lung-on-a-chip device for monitoring the electrochemical and mechanical changing at the lung alveolar interface. Their device was also made of PDMS and mimicked the mechanical strain in alveolar by bulging a porous PDMS-based membrane. The advantage of this model is to create a 3D cyclic strain applied in all directions. An array of suspended gels was used instead of a porous membrane as a barrier between airway epithelial cells and airway smooth muscle cells in another work. There, the gel was a mixture of type I collagen and Matrigel which could enhance cell adhesion and growth. Xu et al.introduced an exciting approach to combine other organs in a lung-on-a-chip for studying lung cancer metastasis. This device had an upstream, lung, and three downstream organs that may be affected by lung cancer metastasis. These are only some of the works done in the field of lung-on-a-chip and more papers with details can be found in a recent review paper published in Biomicrofluics.

In the last few years, several organ-on-a-chip start-up companies have been founded, aiming to introduce a new platform for the drug discovery industry. Some of these companies offer devices at “body-on-chip” scales. Hesperos, Tissuse, CnBio, DRAPER are examples of these companies that are trying to provide a more complex model for assessing the pharmacokinetics and pharmacodynamic of new drugs on the human body than the conventional in vitro cell culture. Indeed, these models cannot completely mimic the complexity of the human body in the absorption, distributions, metabolism, and elimination of these new drugs but they are at least one step ahead of the available models. Some of the companies took another approach to construct an organ-on-a-chip that has a simulated tissue interface. Dr. Donald Ingber at Wyss Institute is one of the pioneers in this area and the founder of Emulate Inc. They offer a range of organ-on-a-chips including lung-on-a-chip or airway-on-a-chip. The strength of their technology is the ease of use for users as they developed several control modules. These modules can be used to control and monitor individual chips — another lung-on-a-chip company, Alveolix, founded by Dr. Olivier Guenatfor simulating the alveolar barrier. Not having vascular perfusion may be considered as a weakness, but it simplified the operation of the chip and would enable the users to operate multiple chips at the same time. Emulate and Alveolix are the most well-known companies in the area of lung-on-a-chip but there are more companies in the field. NORTIS, Quorum Technologies (Artery-on-a-chip Vessel), MIMETAS, SYnVIVO, 4DESIGN BIOSCIENCES, and AIM BIOTECH are good examples of these companies which have focused their efforts to bring some of the complexity of the human tissues to their in vitro models.

Regardless of all these advancements, there are still some issues that need to be addressed to introduce a more biomimetic model. For instance, the cell-to-liquid ratio should be improved to avoid the dilution of secreted proteins, metabolites, and factors. Fabricating smaller channels or growing cells in 3D may help to solve this problem. The thickness of membranes is typically around 10 µm which is much thicker than the basement membrane of blood vessels (300 – 400 nm). Molding hydrogel with integrated microchannels or 3D printing of hydrogels is a possible solution to tackle this challenge. The current models for alveolar can be improved by recapitulating the 3D shape of lungs air sacs. Instead of stretching the membrane that separates apical and basal sides, the membrane can be inflated by flowing air to the air channels.

To sum up, lung-on-a-chip and airway-on-a-chip are only one piece of the puzzle for developing a reliable platform in a drug discovery journey. As the future organs-on-a-chip or body-on-a-chip is supposed to be a comprehensive platform for testing a new drug, multiple chips should be combined to cover all aspects of human physiology. An interesting approach is to design a chip that mimics the human lungs at a smaller scale while having airways, bronchioles, and alveoli at the same time and show how various epithelial cells may interact with each other or which type is more prone to a specific disease or a drug. Realizing such a design can be very challenging based on the current technologies but 3D printing and bioprinting can bridge this gas and facilitate the progress toward a 3D lung model.

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Mohammadhossein Dabaghi

Mohammadhossein Dabaghi is currently a postdoctoral fellow at the department of medicine at McMaster University. He completed his Ph.D. at McMaster University majoring in Biomedical Engineering and worked on new microfabrication technologies to develop an Artificial Placenta device for preterm neonates with respiratory failure. He earned a master’s degree in Biomedical/Chemical Engineering (2014), and a BS in Chemical Engineering from Sharif University of Technology, Iran (2012). His research interests focus on lung-on-a-chips, biosensors for organ-on-a-chips, and bioprinting.

• https://www.researchgate.net/profile/Mohammadhossein_Dabaghi2

The post The Promise of Lung-on-a-Chip Devices appeared first on The MicroFluidic Circle.

In general, PDMS belongs to the silicone family with some unique features which have turned PDMS to the most desirable material in microfluidics. Transparency, biocompatibility, flexibility (Young’s elastic modulus of ~ 1 – 3 MPa), high gas-permeability, low dielectric constant, low surface tension, and low solubility are some of PDMS characteristics.

The PDMS strength stems from its capability in soft lithography. First, PDMS base monomer is thoroughly mixed by its curing agent, then this prepolymer is degassed to remove all air bubbles, and finally, it is ready to be poured on the mold. The mold can be made by conventional methods such as photolithography or newer technologies such as 3D printing. Regardless of the mold type, PDMS can replicate features on the mold from macroscale to nanoscale. After curing and peeling off the PDMS replica from the mold, the PDMS part should be sealed by a flat surface. Here, another advantage of PDMS fabrication comes to play. There are a variety of techniques to seal a PDMS replica such as conformal contact, physical bonding, vacuum bonding, oxygen plasma bonding, corona surface activation, flame bonding, wet-bonding, adhesive bonding, and to name a few. Some of these methods are reversible, and some are irreversible which can be chosen based on applications. In most of these techniques, there is no need to use a chemical or solvent to achieve the sealing. The chemical- or solvent-free bonding of PDMS to PDMS substrates or other substrates eliminates any chance of chemical contamination to tested samples.

PDMS soft lithography allows researchers to fabricate devices with multilayers of PDMS. This process is called “sandwiching,” meaning that several layers of PDMS replica can be stacked on top of each other to build a more complex geometry. In sandwiching, other components such as membranes (porous or non-porous) can be added between layers to create the desired device. There are different ways to bond these membranes to PDMS: (1) membranes first can be treated by a silane molecule (such as 3-aminopropyltriethoxysilane), and then both treated membrane and PDMS are exposed to oxygen plasma and be bonded, (2) some adhesive (double-sided tapes or adhesives) may be used to attach membranes to PDMS if no high pressure is sought, and (3) silicon dioxide (SiO₂) can be coated on a membrane by sputtering and be bonded to PDMS using oxygen plasma.

There are a lot of strategies to tailor the bulk or surface properties of PDMS. This allows researchers to modify their PDMS devices based on their needs. For tailoring bulk properties of PDMS, these methods can be used: (1) changing the ratio of base and curing agent, (2) playing with curing conditions (temperature and time), (3) adding other molecules to PDMS, and (4) adding fillers such as SiO₂ to PDMS. Playing with these parameters can impact the bulk properties of PDMS such as elasticity, transparency, photothermal effect, refractive index, electrical conductivity, and to name a few. The review paper which was written by P. Wolf et al., comprehensively reviews all different techniques for PDMS bulk or surface modification. In most scenarios, surface modification of PDMS is the area of interest in microfluidics. The most common way is to use oxygen or air plasma, corona discharge, and ultraviolet light or ozone exposure to introduce silanol (Si-OH) groups to the surface of PDMS at the cost of methyl groups (Si–CH₃). This results in forming a thin layer of few nanometers on the surface (it may contain some cracks as well) and decreasing water contact angle to lower than 5°. When a long-lasting surface modification strategy is desired, the surface of PDMS can be tailored by polydopamine (PDA), various polyethylene glycol (PEG) derivatives, or silane-based molecules. Hydrophobicity or hydrophilicity of PDMS surfaces can be easily rendered using one of these methods. There is a broad range of strategies and techniques to tailor PDMS properties in the literature which can be considered as one of the advantages of PDMS over other materials for microfabrication.

Tailoring the bulk properties of PDMS to improve its conductivity has become a topic of research in flexible and stretchable electronics. Different fillers such as carbon nanotubes, graphite, silver particles, nanowires, and gold nanotubes have been added to PDMS to reach higher conductivity. As PDA is conductive, it has been coated on PDMS to integrate electrochemical sensors for various applications.

With the origin of organ-on-a-chips and microfluidic cell culture platforms, PDMS has gained more attractions from researchers from other fields. As a result, the surface of PDMS has been coated with PDA, gelatine, collagen, or fibronectin to enhance the adhesion, proliferation, and the growth of cells inside a chip. Besides, the surface of PDMS first can be modified by PEG or a silane molecule or PDAas linker, and then a bioactive molecule can be attached to the linker for achieving specific applications such as differentiation or detection of a secreted biomolecule by cells.

PDMS is permeable to gases such as oxygen and carbon dioxide. This led to the origin of microfluidics blood oxygenators in which gas exchange between blood and air is needed.PDMS membranes can be easily fabricated by spin-coating wet PDMS on a flat substrate such as a wafer. The properties of the PDMS membrane can be tuned by changing the speed of the spinner, curing agent ratio, and curing temperature. PDMS membranes as thin as ~ 1 µm can be produced without observing pinhole defects. When the permeability should be avoided, the thickness of PDMS can be increased over 1 mm.

In a nutshell, PDMS has shown its potential in the realization of so many applications in microfluidics. Microfabrication with PDMS enables us to fabricate devices with simple designs to devices with very complex features. PDMS has been extensively studied, and this provides fantastic resources for everyone to optimize PDMS properties based on their needs.

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Mohammadhossein Dabaghi

Mohammadhossein Dabaghi is currently a postdoctoral fellow at the department of medicine at McMaster University. He completed his Ph.D. at McMaster University majoring in Biomedical Engineering and worked on new microfabrication technologies to develop an Artificial Placenta device for preterm neonates with respiratory failure. He earned a master’s degree in Biomedical/Chemical Engineering (2014), and a BS in Chemical Engineering from Sharif University of Technology, Iran (2012). His research interests focus on lung-on-a-chips, biosensors for organ-on-a-chips, and bioprinting.

• https://www.researchgate.net/profile/Mohammadhossein_Dabaghi2

The post PDMS and Its Role in The Realm of Microfluidics appeared first on The MicroFluidic Circle.

Preterm (premature) babies with RDS (respiratory distress syndrome) require additional respiratory supports to let their lungs develop. Usually, mechanical ventilation is the first option to deal with RDS. If mechanical ventilation is not sufficient, another life support device, named extra-corporeal membrane oxygenation (ECMO) machine, will be used. In ECMO, the blood is pumped outside of the body and oxygenated by a gas exchange device called the blood oxygenator. Both methods are invasive and can cause severe long-term side-effect and complications. In mechanical ventilation, positive pressure applied in the lungs can seriously damage the tissue of the lungs. ECMO may be less destructive to the lungs but it has own issues, especially for those babies with extremely low birth-weight. Generally, ECMO devices need a high volume of blood to be primed (for instance around 40 mL for the device such as Quadrox-i Neonatal & Pediatric), however, a one-kilogram preterm baby has only about 100 mL of blood. As a result, donor blood from adults needs to be transfused to keep the whole system operational. Nonetheless, baby blood and adult blood have different physiology in which red blood cells (RBCs) of newborn babies have a higher affinity to oxygen compared to RBCs in adults. Additionally, surgery is required to get access to the large vessels and a pump is needed to keep blood flowing in the circuit which may damage RBCs and lead to hemolysis and thrombus formation.

Artificial Placenta as an Alternative Solution

The artificial placenta (AP) concept refers to a postnatal lung assist device (LAD) which is connected to a premature baby via the umbilical cord and can provide additional gas exchange besides the baby’s lung. The AP concept has been around for a long time back to 1961 when the first attempts were made by Callaghan et. al. However, there is still no commercial AP device in clinics due to inefficiency and incompatibility of available LADs for this purpose. An ideal AP device should be operable by a baby’s heart instead of an external pump while provides sufficient gas exchange in room air. Also, such a device should have very low priming volume to eliminate any need for donor blood transfusion and be filled only with saline solution. In addition, the volume of needed saline should not surpass 10 percent of the total blood of the baby’s body to avoid blood dilution. The first and main job of an AP device is to compensate for the loss in gas exchange caused by under-developed lungs. Also, some effort was made to expand AP application more and provide nutrition to premature babies and remove the metabolic waste at the same time.

Conventional blood oxygenators (used for ECMO) which are usually made from hollow fiber membranes are not suitable for artificial placenta application as they have high priming volume and need a pump for operation. A hollow fiber membrane (HFM) device is made by twisting bundles of hollow porous tubes (fibers) in a plastic-made shell. Usually, oxygen is introduced to these fibers and blood flows around these porous fibers where oxygen transfers to blood and carbon dioxide is released from the blood. Because of the nature of the hollow fiber bundle, the blood experience a non-physiological travel which may damage RBCs and initiate clotting. Also, there is a great potential of blood leakage through pores or contamination of the blood. An ECMO device is connected to the body in venous-artery or venous-venous configuration where a pump is needed to remove blood from the body and send it back to the ECMO device and vice versa.

It should be noted that an AP device is not going to operate in the same way that an ECMO machine works. In fact, an AP device is connected to a baby’s body in parallel to the blood circulatory system wherein the inlet of the device is attached to an artery and the outlet to a vein. This particular type of arterio-venous connection causes the blood entering the device has already been traveled through the baby’s lungs and be partially oxygenated. Under-developed lungs of a premature baby cannot fully saturate the blood with oxygen, whereby the oxygen saturation level may only be 10%–30% below full saturation. Therefore, an AP device only needs to raise the oxygen saturation level from ~ 70 % to 100 % which is higher than the amount of oxygen saturation level entering an ECMO device.

The Solution

Over the last decade, researchers have been using microfabrication technologies to overcome the shortcoming of HFMs and trying to introduce the next generation of blood oxygenators. Microfluidic blood oxygenators have shown promising improvement in gas exchange compared to convention ECMO devices over the last decade due to mimicking nature lungs by reducing the size of blood channels as small as capillaries in the lungs. Additionally, microfluidic blood oxygenators provide a more physiological flow path to the blood which is critical in designing an AP device for the long-term application.

Recently microfluidic blood oxygenators have been specifically designed to address the needs of the neonates and have been developed with low priming volume, low-pressure drop to operate pumpless, and capable of working in the ambient air. Rochow et. al.used microfluidic blood oxygenators to construct such a LAD for premature babies suffering from RDS. Their device had a low priming volume of 4.8 mL and could be successfully connected to the umbilical vessels of newborn piglets and perfused solely by the pressure differential provided by the animals’ heart. Although the gas exchange of this LAD was not sufficient enough but they showed that a pumpless LAD with low priming volume was not a fantasy anymore!

Challenges

Truly, the future belongs to blood oxygenators advancing microtechnologies owing to their unique design so much similar to nature lungs. However, they need to overcome some challenges to be able to compete with HFMs and replace them in the markets. The main challenges are listed below:

1. Scaling up: the main and most important limitation of all microfluidic blood oxygenators is their small size which constrains their operating blood flow rate below 10 mL/min. However, an AP device may need to support a higher blood flow rate up to 60 mL/min. Different strategies have been used to address this issue such as increasing the size of the device or connecting several devices together to increase the total capacity. It seems like a combination of both should be used to reach the desired gas exchange.

2. Hemocompatibility: although different approaches such as coating the surface of PDMS by antithrombin-heparin (ATH) or polyethylene glycol (PEG)have been used to improve the hemocompatibility of the blood-contact surfaces, more investigation is required to evaluate this coating reliability for a long-term condition as needed for an AP device.

3. Access to large bore vessels: an AP device is supposed to be connected to a premature baby via the umbilical cord. The main challenge is the constriction of the umbilical vessels right away after birth. As a result, a catheter is required to be inserted into the vessels and keep the vessels open and let blood flow with minimum resistance to blood flow. As Peng al. explained, a special type of catheter should be used for getting access through the umbilical cord to be able to minimize the resistance to blood flow and provide the maximum possible blood flow rate for an AP device.

Enjoyed this article? Don’t forget to share.

Mohammadhossein Dabaghi

Mohammadhossein Dabaghi is currently a postdoctoral fellow at the department of medicine at McMaster University. He completed his Ph.D. at McMaster University majoring in Biomedical Engineering and worked on new microfabrication technologies to develop an Artificial Placenta device for preterm neonates with respiratory failure. He earned a master’s degree in Biomedical/Chemical Engineering (2014), and a BS in Chemical Engineering from Sharif University of Technology, Iran (2012). His research interests focus on lung-on-a-chips, biosensors for organ-on-a-chips, and bioprinting.

• https://www.researchgate.net/profile/Mohammadhossein_Dabaghi2

The post Breathing Like in the Mother’s Womb appeared first on The MicroFluidic Circle.