In traditional drug screening approaches, organ-specific effects and the toxicity of a drug are examined using a homogeneous population of specific cells. However, human tissues and organs are not made of a homogeneous cell population. The complexity of human organs is based on intricate interactions between various specialized cell types arranged in precise geometries and interacting with specific microenvironments. These interactions often occur at well-defined tissue interfaces, enabling organ function.
One of the first papers detailing the use of organized cell cultures to study disease was published by Andre Kleber in 1991, reporting the construction of a ventricular myocardium through the patterned growth of cells in vitro, which enabled the first biophysical explanation of conduction block in the heart. The field of biomicrofluidics exploded in the late 1990s with the introduction of poly(dimethylsiloxane) (PDMS), which is an optically transparent, soft elastomer ideal for biological applications on the small scale. The concept of mimicking the organ-level function of human physiology or disease using cells inside a microfluidic chip was first published in 2004, when Michael Shuler and colleagues first demonstrated a system that captured the systemic interaction between lung and liver on a one square inch silicon chip. A range of microfluidic devices have since been developed, mimicking diverse biological functions by culturing cells from blood vessels, muscles, bones, airways, liver, brain, gut and kidney. In 2010, the term organ-on-a-chip was invented by Donald Ingber, who developed a microfluidic chip to capture organ-level functions of the human lung.
This project was started as an attempt to overcome the limitations of the existing PDMS-based Organ-on-Chip systems, some of which are:
- PDMS has higher stiffness (~1000 kPA) compared to extracellular matrix or ECM (0.5-2 kPa).
- Higher stiffness of PDMS leads to non-physiological growth of cells. Furthermore, hydrophobicity of PDMS also leads to poor adhesion of primary mammalian cells.
- Cells cannot remodel their environment as an adaptive response, unlike in-vivo.
- PDMS membranes have a regular array of micropores in many Organ-on-Chip devices, which may exaggerate the chemoattractant signaling, and hence inflammation.
These limitations could potentially be overcome by making Organ-on-Chip devices using hydrogels such as collagen, alginate and gelatin. They are usually as transparent as PDMS and have stiffness of ~1-5 kPA which is more physiological. Therefore, in this project, a detailed investigation of challenges involved in 3D printing of hydrogels was done. The fabrication protocol for the following design was inspired by Lewis Lab‘s work on Bioprinting of 3D Convoluted Renal Proximal Tubules on Perfusable Chips.
The idea underlying the design (Fig. 1) was to have two closely spaced perfusable channels embedded within hydrogel to have effective trans-hydrogel cytokine signalling, and consequently immune response. 3D Printed sacrificial Pulronic F127 channels were embedded in Gelatin Methacryol (GelMA). Channel geometry and printing parameters were optimised to minimise the separation between channels and a reproducible spacing of 50 um was achieved. Further optimization needs to be done in order to achieve confluency within the channels (Fig. 6).
This work was done in McKinney Lab, EPFL School of Life Sciences, Switzerland under the supervision of Prof. John McKinney and Dr. Vivek V. Thacker.
- Zhang, B., Korolj, A., Lai, B.F.L. et al. Advances in organ-on-a-chip engineering. Nat Rev Mater 3, 257–278 (2018).
- Homan, K., Kolesky, D., Skylar-Scott, M. et al. Bioprinting of 3D Convoluted Renal Proximal Tubules on Perfusable Chips. Sci Rep 6, 34845 (2016).