How to reproduce active biomimetic stimulation in vitro?

Active mechanical stimulation

Compressive stress

Specific tissues such as heart, blood vessels, and bone cells are exposed to compression^1–3. A great example of the importance of compression is cartilage regeneration. It is proven that mechanical stimulation of intervertebral disc cells (that mostly originates from compression) acts as an important stimulus for increasing cartilage matrix anabolism.

In vitro investigations demonstrated that without any physical stimulation, chondrocytes change their phenotype during proliferation and lose their capability to produce the typical extracellular matrix for cartilage⁴.

Mechanical stimulation is hence a prerequisite for in-vitro investigations on cartilage and spinal diseases, and compression has well been studied in bone and cartilage tissue engineering^4,6. 

Fluigent presented a cartilage on a chip model (see figure) to elucidating how chondrocytes react to mechanical stimuli, allowing to understand processes triggering cartilage diseases like osteoarthritis¹¹. This organ on a chip platform (consisting of a mechanical stimulation unit, a PDMS membrane, a 3D cell culture chamber, and a perfusion channel) allowed performing mechanical stimulation on cells, while creating dynamic culture conditions. More information can be found on the application note webpage.

In the microenvironment, compression imposed on a cell generally originates from cell-cell interaction or cell-ECM interactions. Organ on a chip devices can be used to mimic compression by applying a pressure on specific areas of the chip, thus allowing for compressing cells or tissues in a precise manner. The devices usually integrate flexible membranes or diaphragms which are bent upon compression.

Membrane compression is mostly performed by applying air pressure on top of the membrane. Evaluating the pressure applied on the membrane is easy, but determining stress directly exerted on cells or tissues is however more difficult as stress distribution will depend on the design of the organ on a chip device.

Stress estimations are usually performed using finite elements analysis modeling, where one can implement channel dimensions, material properties, and applied pressures^7,8.

Several examples of compressive stress in organ on a chip can be found for tissue engineering and regenerative medicine. A good example is the organ on a chip model developed by Sim et al⁹. Traditional methods like hydrostatic pressure chambers generally use a high amount of compression load¹⁰, limiting the studies to high compressive stresses. In addition, such methods often require large number of cells, samples, or fluid volumes. To tackle these limitations, an organ on a chip device containing microscale cell culture chambers separated from an air-pressure chamber was developed, allowing to apply micromechanical stimulations⁹.

The authors compared human mesenchymal stem cell proliferation using cyclic compressive stress at moderate levels (1 to 5 kPa) and at static conditions. They found that the cell proliferation rate of the stimulation group was about 1.5 higher than the control group after one week, revealing that cyclic compressive stress at moderate levels can enhance the proliferation of human mesenchymal stem cells.

Stretch and strain

Pulsatile flow is mainly observed in arteries. Pulsations are smoothed by the muscle of the arterial walls. Pulsations are a direct consequence of heartbeats. The heart acts as a reciprocating pump that drives blood directly into the aorta. At each stroke, the flow reaches a peak (systole), then diminishing to a low (diastole) until the next stroke⁶. This produces pulsatile flow at each stroke instead of a continuous flow. Although pulsatile, flow remains laminar, the velocity flow profile varies as a function of time⁶. In fact, a typical flow rate curve of the artery for one heartbeat cycle displays two local maxima and a minimum, with positive and negative flow rate values (figure 2 a)⁷. As a consequence, the flow direction and the amplitude of the velocity flow profile will vary as a function of time (figure 2 b). This, of course, has an impact on the shear stress (or shear flow), as it is derived from flow velocity. More information on flow velocity patterns and subsequent shear flow can be found in the literature⁶. Pulsatile flow is usually performed in blood vessel-on-chip models to simulate the actual pulsatile blood flow in human circulation⁸.

In addition to flow-induced shear stress and compressive stress, tensile force applied to the cellular microenvironment leads to strain and stress to the tissue. The lung is a good example of this phenomena. During inspiration, intrapleural pressure decreases, causing the alveoli to expand. This pulls air into the lungs, resulting in stretching of the alveolar epithelium and the endothelium in adjacent capillaries¹². This stretching results in mechanical strains induced on cells. It has already been demonstrated that tensile stress induced by stretching affects cellular membrane integrity¹³, cell shape, spreading, and proliferation¹⁴. Stretching is usually performed for gut on a chip, heart on a chip or muscle on a chip models as these organs experience stretching or strain in vivo^13–15. Stretching in organ on a chip devices is usually performed using pressure or vacuum^13,16–20. The chips usually consist of a central channel for cell culture and perfusion, and pressurized side channels (below, or around the central channel, see figure). The channels are made of flexible material, allowing for elastic distortion of the membrane when applying positive or negative pressure. Upon bending, cells or tissues are stretched and experience a tensile stress.

It is possible to estimate the strain and stress applied on the device and cells using theoretical models derived from continuum mechanics, that are computed in finite element models. The models are relatively complex and can differ from one system to another (bi-axial or circumferential strain, gel or cell culture …). Excellent models description and derivation can be found in the literature^{13,17,18,20–23}. Briefly, the authors used nonlinear or hyperelastic material models that connect stresses to strains, using tensors such as the Cauchy-Green strain tensor and the Piola-Kirchhoff stress tensor. The models are generally implemented in finite element software for validation.

A great example of how organ on a chip technology with stretch-induced cell stress allowed scientific progress can be found in a gut on a chip model developed by a joint group of researchers from Harvard university and the MIT. Human intestinal inflammatory diseases such as Crohn’s disease are believed to be caused by a combination of many factors such as complex interactions between gut microbiome²⁴, intestinal mucosa²⁵, and peristalsis suppression²⁵ (strongly associated with inflammation and intestinal bacterial overgrowth).  It is however not possible to study the intrinsic effect of each factor in vivo as they are inextricably linked. Traditional and in vitro models (i.e cell culture in flasks or petri dishes) are also incompatible with such studies as they poorly reproduce the pathophysiology of human inflammatory bowel disease. In such models, peristaltic motion which is a strong driver of normal cell differentiation is missing. To address the above limitations, Kim et al. developed a human gut on a chip device consisting of two microfluidic channels separated by a porous flexible membrane coated with extracellular matrix, and two side vacuum chambers for simulating peristaltic motions and subsequent cell stretching¹⁵. The devices allowed them to study the factors of intestinal inflammatory diseases individually. In particular, they used the chip to analyze the effect of cyclic stretching (10% in cell strain, 0.15 Hz in frequency) on the proliferation of GFPEC bacteria cultured under flow. They observed that bacterial cell densities more than doubled within a day when cultured without cyclic stretching.

This discovery suggested that cessation of epithelial distortion can trigger bacterial overgrowth, and refuted the previous hypothesis that bacterial overgrowth originated from fluid flow²⁶. This new organ on a chip model allowed one to analyze several factors in a controlled manner, which was not possible using existing in vitro systems or animal models, hence allowing to gain new insights into gut pathophysiology.

Shear from constant laminar flow

Interstitial fluid flow is the movement of fluid through the extracellular matrix of tissues, where cells such as fibroblasts, immune tissue cells, and adipocytes can be found^1,9. Fluid flow carries large proteins through the interstitium and mechanically stimulates interstitial cells. Several studies have demonstrated that shear flow induced by interstitial fluid was crucial for cellular activities, as such flows induced physiological responses from cells^10–14, such as cell differentiation. Interstitial fluid typically flows at a lower velocity compared to blood flow within vessels because of the high flow resistance of the extracellular matrix. Flow velocity profile and subsequent shear flow is also more difficult to define due to the complex architecture of the extracellular matrix and as the fluid moves around the cell-matrix interface in all directions. Some studies have analyzed in-depth the shear stress with such architectures using numerical simulations.

A 3D cell culture microfluidic device was developed to provide new insight on how interstitial flow affects breast cancer cell invasion¹⁵. Specifically, the authors found that compared to static flow conditions, interstitial flow increased the number of migratory cells, as well as their migratory speed. These observations demonstrate how important it is to consider interstitial flows in tumor models, as they affect tumor cell invasion and invasion direction.

References

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