a,b | With biological compatibility and tunable mechanical properties, poly(dimethylsiloxane) (PDMS) is fabricated using soft lithography to develop an elaborate microstructure323 (part a) or 3D rapid prototyping by laser processing47 (part b). c | Integrated device fabrication with gold- or platinum-deposited substrate such as an electrode array324. d | For large-scale experiments such as high-throughput screening, plastic-based injection moulding could be selected50,51. e,f | For rapid prototyping with a high degree of design freedom, resin-based (part e) or hydrogel-based (part f) 3D printing is available325,326. HCS, high-content screening.
Media perfusion is a hallmark of OoC devices, serving as a circulatory system mimic that maintains a concentration gradient for nutrient and waste convective transport. To drive media perfusion through the OoC device, various pumps have been adapted for OoC applications (Supplementary Fig. 2). These include conventional syringe pumps75,76, microvalve-driven actuator pumps77,78 and peristaltic pumps38,79. Alternatively, various groups have also adopted hydrostatic pressure-based pump-free systems to drive perfusion80,81,82. The choice of pumps will greatly depend on whether the culture runs on a one-pass type of perfusion flow or a recirculatory flow configuration. Conventional syringe pumps and pump-free gravity-driven flows typically support one-pass perfusion flows, whereas other pump variants such as peristaltic pumps are amenable to driving recirculatory flow. Recently, however, we have seen the emergence of OoC devices with gravity-driven recirculatory flows, which removes the need for physical pumps, hence minimizing overall system complexity while still mimicking soluble factor crosstalk between multiple organ compartments82,83,84,85,86. The choice for either type of flow and for the flow rate (volume per time) depends on the needs of the organ model and the in vivo situation, including the appropriate level of shear stress induced by flow over a tissue surface.
A first multiple-organ system on a chip was designed to mimic systemic distribution of, and toxicity response to, the compound naphthalene106. The device contained a combination of four compartments representing the lung, liver, fat and other tissues. The dimensions of the tissue compartments and microchannels connecting them were specifically designed to mimic systemic blood flow distribution and residence times in the respective organs; after passing through the lung, fluid flow was volumetrically distributed to the fat (9%), liver (25%) and other tissue (66%) compartments. The study was focused on mimicking the toxicity of naphthalene metabolites, generated by the liver, on lung cells. The different compartments, therefore, had essential interconnections in the form of microchannels that allowed the distribution and metabolism of these compounds (conceptualization and design). The device was assembled from etched silicon and machined PMMA, which yielded very well-defined channels. These materials allowed for optical inspection of the compartments on the chip (material selection and fabrication). The lung and the liver compartments contained rat cell lines, whereas the fat and other tissue compartments remained cell-free, thereby not participating in the biological response but serving to better emulate the proportion of drugs distributed to the biologically active liver and lung compartments (selection of biological elements). Peristaltic pumps were used to recirculate media through the different organ compartments (supporting life inside the device). The distribution, metabolism and toxicology of compounds could be studied with this device. The introduction of a flow component into conventionally static cell culture models allowed the study of inter-organ interactions, which would not have been possible in a non-OoC format. However, extrapolation from this miniaturized rat-based system to humans remained difficult. Today, it has become possible to use human organoids or spheroids, or even simpler cell lines of human origin, in this type of experiment to eliminate the need to take species differences into account.
A different approach for a multi-OoC for drug metabolism was developed later, aimed at developing a physiologically relevant multi-organ system for drug discovery107 (Fig. 3B). This system included four, seven or ten organ models, which were coupled with integrated fluidic circuitry to study organ interactions (conceptualization and design). This integrated device was fabricated with easier to process polysulfone and polyurethane, with integrated air pressure-controlled valves as built-in pumps to recirculate media in a more controllable fashion (material selection and fabrication). The organs modelled were of human origin (primary cells, from cell lines and iPSC-derived cells) and included the kidney, muscle, liver and gut, amongst others (selection of biological elements). The pumping system was incorporated into the final device, requiring only external air pressure to regulate the valves by software (supporting life inside the device). The drug diclofenac was administered to the gut section of this system, where it was absorbed and, subsequently, distributed among the other organ models. The complexity that this system allowed was necessary to answer a more complex scientific question, but was met with an increase in the complexity of the peripheral system, with various pieces of supporting equipment required (including a 36-channel pressure/vacuum controller). This added complexity leads to additional costs for setting up the system, and a large laboratory footprint for a miniaturized system.
Many OoC system designs and operations are focused on trying to reproduce physiological tissue architectures and functions, and therefore less emphasis is placed on testing assays and read-outs. However, for OoCs to be routinely deployed for standardized testing of drugs and chemicals, such as those stipulated in OECD testing guidelines, the compatibility of the designed OoCs with common laboratory instruments for data acquisition and downstream analyses is an important practical consideration. If OoC devices can be designed to have reagent interfacing and cell imaging layouts that are similar to those of standard culture plasticware and glassware, the technology will be more approachable for biological end users as the workflow and associated hardware and software for analyses would be familiar and mostly already present in the laboratory. An example involves fitting 20 microfluidic devices in a 96-well PS plate and using a plate reader to quantify luminescence in each device270.
To support broader use in biological research and drug development, standardized OoCs need to become as available as cell culture plates, implying that they need to be reliably manufactured on an industrial scale269. Approaches such as injection moulding, long established in the manufacture of high volumes of plastic products at low cost, are attractive for OoC manufacture where their context of use demands high throughput50,51. As OoCs often consist of several layers, the ability to manufacture many plastic chips may need to be coupled to parallel developments in the automation of chip alignment, assembly and bonding. With increased awareness and emphasis on manufacturability, we are also seeing more OoC systems being fabricated from new materials using additive manufacturing techniques such as 3D printing291. These techniques are constantly being improved with respect to production speed, and provide a largely assembly-free approach to device construction. They are thus a promising alternative for high-volume OoC production. The other major advantage of additive manufacturing is the ability to rapidly prototype new OoC designs. This can have a profound influence on how OoC designers develop their systems, as iterative refinement can be performed far more rapidly and effectively. 2b1af7f3a8