1. Introduction
Synthetic polymers with a -R2-Si-O unit are known as silicones, where the -Si-O repeating unit is called a siloxane. Both thermal and chemical stability are given by the strength of the -Si-O bond [1]. Poly(dimethylsiloxane) (PDMS) is the most popular silicone elastomer to fabricate microfluidic devices, with the Sylgard 184® from DOW-Corning, USA, being the most frequently used to study fluids at both the micro- and macro levels [2,3,4,5]. PDMS is an elastomer that has high flexibility and elasticity. In addition, the flexibility of the siloxane backbone allows the chains to easily arrange and rearrange themselves, and can be changed by adjusting the ratio of the base polymer (monomer) to the curing agent (crosslinker) [1,3,6]. Many researchers have been using PDMS in a variety of engineering applications, including mechanics, electronics, and biomedicine, due to their outstanding properties [1,7,8,9,10,11,12]. Thus, PDMS has been widely used in a range of applications, including micropumps [13], microvalves [14], microneedles [15,16], microreactors [17], optical systems [18], biomodels [5,19,20], blood analogues [21], and microfluidics [12,22,23,24,25].
2. PDMS Properties
In the early days of microfluidic research, the majority of the phenomena, discoveries and visualizations were performed in glass microchannels [26]. One of the most relevant blood flow studies was the work carried out by Goldsmith in 1971, where he investigated the trajectories, velocity profiles, interactions, and deformability of human red blood cells (RBCs) in glass microchannels [27,28]. However, PDMS, due its unique properties such as simple fabrication, excellent optical transparency, biocompatibility, gas permeability, effective adhesion to itself and to the glass, and flexibility, is nowadays the most widely used material to fabricate microfluidic devices and consequently to study flow phenomena happening at the micro-scale level. In addition, recent microfluidic applications usually require flexible materials since the walls of the device may need to be elastic, i. e., microchannel walls, in several applications, should return to their original shape after removing the stress induced by the fluid flow [3,12]. Figure 1a shows a schematic diagram of the most usual polymers used in microfluidics, whereas Figure 1b highlights the predominance of PDMS in microfluidic devices compared to Polystyrene (PS) and Polymethylmethacrylate (PMMA). The popularity of PDMS can be attributed to its flexibility and adjustable elasticity, which can be easily modified by combining the weight ratio of the base polymer and the curing agent.
PDMS is an elastomer with long chains that has a modulus of elasticity of 1–3 MPa. When compared to about 50GPa of glass, PDMS can be considered as a material with high flexibility and elasticity [1,12,29,30]. PDMS, besides being chemically inert and thermally stable, can replicate submicron features of molds to develop both simple and complex microstructures [31,32,33]. Furthermore, it can be easily bonded to different kinds of materials by using plasma surface treatment or other alternative methods [34]. Hence, most current microfluidic studies use microdevices made of PDMS instead of using materials like silicon and glass [1,12,35]. The excellent optical transparency of PDMS makes it possible to perform both qualitative and quantitative real-time observations of microflow phenomena. Optical access and high-speed image acquisition is essential to understand physical and biological phenomena happening in microfluidic devices [36,37,38,39,40,41]. Another extremely important feature of PDMS is its biocompatibility, which means that it is compatible with biologic tissues and in this way suitable to develop organs-on-chip (OoC) [31,42,43]. However, from a biomedical perspective, the primary drawback of PDMS is its hydrophobic nature, which is not beneficial for applications that need surface adhesion such as cell culture [44,45]. In addition, hydrophobic surfaces can allow liquid molecules, gases, fluorescent dyes, and other kinds of chemicals to adsorb and absorb on microchannel walls, which can be a disadvantage for several biomedical applications [31]. The natural hydrophobic surface of PDMS poses several challenges for fluid handling applications. Hence, several studies have been performed to improve PDMS wettability and in this way to allow efficient liquid flow without compromising the structure or functionality of the device. With increasing research on surface treatments to enhance these characteristics, various techniques have emerged, including oxygen plasma treatment [44], UV irradiation [46], and surfactant-based modifications [47]. These methods aim not only to optimize flow in microfluidic devices, but also to improve the separation efficiency of cells and particles of different sizes and reduce issues such as bubble formation and particle aggregation [3,48,49,50]. In Table 1, several relevant properties of PDMS for fluid applications are listed.
3. Recent Advances of PDMS Applications in Engineering
PDMS has been widely used in biomicrofluidics and in other fields in engineering such as in heat transfer and face masks. Recently, PDMS has played a crucial role in the progress of OoC systems, which are bioengineered microdevices that mimic key functions of human organs and tissues at a micro-scale level. These platforms are being used to study drug interactions, disease progression, and cellular responses in real-time [31,42,43]. Another significant advance in PDMS applications is the development of blood flow models, also known as phantoms or simply as biomodels. These devices are mostly fabricated at a macro-scale level and are commonly used to replicate the flow of blood phenomena and other fluids within the human body. These biomodels allow researchers to study the behavior of diseases such as aneurysms and stenosis under controlled laboratory conditions [5,19,65]. Figure 2 shows several relevant PDMS applications in different fields of engineering.
3.1. Application of PDMS in Biomicrofluidics
Since the beginning of the twenty-first century, the biomedical field has paid attention to PDMS due to its notable advantages. The primary benefits and drawbacks of PDMS and other materials that are frequently used in biomedical microdevices, biomodels, and OoC systems are displayed in Table 2.
Although PDMS has several drawbacks, this polymer has become the most popular material used to fabricate microfluidic devices, in vitro biomodels, and OoC platforms due to its optical transparency, gas permeability, low-cost fabrication, variable mechanical properties, and biocompatibility. However, PDMS has limitations, such as hydrophobicity, absorption of hydrophobic molecules, difficulty in mass production, and attenuation of acoustic waves, which restrict its use in acoustofluidic systems. PDMS does promote the attenuation of acoustic waves due to its viscoelastic isotropic material [58]. Although there have been several successful applications of PDMS in acoustofluidics [58,69,70,71,72,73], harder materials such as glass, silicon, and specific thermoplastics are preferred due to their lower attenuation and better acoustic coupling. An example was the work performed by Qu and Qiu [74], where they developed a microfluidic device to study the effect of an acoustic field on the bubble dynamic phenomena. Although the top part of the device was made of PDMS, the bottom part was made of glass due to its lower attenuation and better acoustic coupling.
Thermoplastics, such as PMMA, are also popular for the fabrication of microfluidic devices and they can be used for mass production via injection molding. While they offer optical transparency, they lack gas permeability and flexibility, making them less suitable for dynamic biological applications compared to PDMS. The current 3D printing resins enable rapid prototyping and the fabrication of complex geometries, providing design flexibility that surpasses PDMS. However, these materials face several challenges such as low optical transparency, limited gas permeability, and rough surfaces that affect fluid flow and cell attachment, reducing their employment for imaging-based biological studies. Emerging 3D printing resins under development offer a promising solution to these limitations, presenting a potential alternative to PDMS for fabricating in vitro biomodels and phantoms at both the milli and macro-scale level [75]. Hydrogels offer distinct advantages, including high biocompatibility, permeability to small molecules, and support for 3D cell cultures within their matrices. Unlike PDMS, hydrogels enable a more biomimetic environment for tissue modeling. However, their weak mechanical strength, degradation over time, and storage challenges (like freezing or drying) limit the applications of hydrogels to produce microfluidic devices and biomodels. Glass is a distinctive material due to its optical transparency, smooth surfaces, and chemical inertness, making it ideal for performing flow visualizations at both the micro- and macro-scale level. Advanced techniques such as nanoimprint lithography and femtosecond laser writing enable glass to be molded into nanometer-resolution geometries, expanding its potential to produce molds with complex geometries [76,77,78,79,80]. However, compared to PDMS, glass remains rigid, fragile, and associated with higher cost constraints, which limits its application in fabricating microfluidic devices and in vitro biomodels. Silicon offers high-precision micro- and nanoscale fabrication, superior thermal stability, and chemical resistance. However, it is costly, opaque, and most of the cases requires clean-room facilities for the fabrication of the devices, unlike the low-cost fabrication of PDMS. Each material shown in Table 2 offers unique strengths for specific applications, but PDMS remains a preferred choice not only to fabricate microfluidic devices and in vitro biomodels, but also for cell-based studies due to its biocompatibility, gas permeability, and cost-efficiency.
Despite the beneficial PDMS properties that have increased its popularity amongst engineers, there are several limitations that need to be overcome. For instance, the hydrophobic surface of PDMS may restrict its use in some biological samples. Because PDMS is hydrophobic, it makes it difficult to wet channel surfaces with liquids and raises microchannel flow resistance [80]. Furthermore, the absorption of molecules on microchannels makes quantitative drug analysis difficult [2,12,53,81]. Several research works have been conducted on changing the PDMS surface to be hydrophilic. Surface modification techniques are commonly used for PDMS surface oxidation in order to enhance microchannel wettability and overcome PDMS hydrophobicity. Surface modification techniques, such as oxygen plasma treatment [44,80,81,82,83], UV treatment [46], and chemical methods [84], have been employed to address PDMS hydrophobicity. Among these methods, oxygen plasma treatment is the most conventional and widely used approach. It generates silica-like surface layers on the PDMS material, effectively increasing hydrophilicity. However, this modification is not permanent, as PDMS gradually recovers its hydrophobicity in a few hours [48]. To overcome this limitation, combining oxygen plasma treatment with additional coatings, such as polyethylene glycol (PEG), has shown potential for prolonging the hydrophilic state [80]. An alternative and viable strategy involves bulk modification through the incorporation of surfactants into the PDMS matrix. Surfactants reduce surface tension, allowing aqueous solutions to spread more quickly over the surface [48]. Consequently, the use of surfactants for surface modification has proven to be a promising approach for achieving long-term hydrophilicity, offering a simple and efficient solution that eliminates the need for complex procedures [85,86].
Long et al. [80] have shown that oxygen plasma treatment followed by PEG coating creates a hydrophilic surface on PDMS. This modification enabled capillary force-driven flow in microfluidic channels, with the Rhodamine B fluid filling the channels in just 13 s (see Figure 3), while untreated PDMS showed no flow even after 60 s. The treated PDMS maintained its hydrophilic properties, supporting fluid flow without external pumping for 420 h. Similarly, Peterson et al. [83] showed that oxidized PDMS coatings on glass–silicon microfluidic devices provided undisturbed flow rates for 14 min, unlike uncoated devices, which experienced a 12% reduction in flow rate, whereas pure PDMS has shown an almost 40% reduction.
Surfactant-based approaches offer a simpler, cost-effective alternative for achieving long-term hydrophilicity. Holczer et al. [87] developed an autonomous capillary-driven system using surfactants for bioanalytical devices. Vilčáková et al. [88] explored CNT-based PDMS composites with surfactants like dodecylbenzene sulphonic acid, cetyltrimethylammonium bromide, and their mixtures. These composites were prepared via mechanical mixing and sonication, achieving a homogeneous dispersion of fillers, improving PDMS wettability. Bulk modifications have also proven effective. Wu and Hjort [89] incorporated Pluronic F127 into PDMS pre-polymer before curing. The Pluronic F127 molecules migrated to the PDMS–water interface, reducing the contact angle from 104° to 63° after water immersion for 24 h. Gonçalves et al. [90] tested Pluronic F127, polyethylene glycol, and polyethylene oxide (PEO) for bulk modification. The 2.5% PEO modification showed superior performance, facilitating fluid flow, reducing cell aggregation, minimizing air bubble trapping, and enhancing sample purity in blood plasma separation.
These studies show that surfactant-based and bulk modifications offer long-term hydrophilicity and improved fluid control in PDMS-based microfluidic devices. This approach offers a simpler and cost-effective alternative for bioanalytical and diagnostic applications.
3.1.1. Application in Microfluidic Devices with Contractions
Photolithography is the most popular method for fabricating microfluidic devices. However, it is more expensive than soft lithography [91]. The fabrication of PDMS microfluidic devices typically involves the following steps:
Mold Preparation: After photomask creation, a master mold is made using photolithography, typically on a silicon wafer with a SU-8 photoresist.
Mixing and Degassing: The PDMS base polymer and curing agent are mixed in a specified ratio (e.g., 10:1) and degassed to remove air bubbles.
Casting: The PDMS mixture is poured into the mold and cured at 60–80 °C.
Demolding: Once cured, the PDMS device is peeled off the mold.
Bonding: Plasma treatment or other surface modification techniques are used to bond PDMS to itself or other materials, such as glass [39].
To reduce costs, alternative methods that avoid clean rooms have been developed for mold and device production [39]. Among the various microfluidic devices, those developed to assess pathological cell behavior have been receiving particular attention [38,92]. A key biomarker to distinguish healthy cells from diseased cells is blood cell deformability [38]. To study this phenomenon, PDMS microfluidic devices have been developed to support disease diagnosis and research, including for cancer [93], diabetes [92], and malaria [94]. Advances in microfabrication [38,39], flow visualization [95], and image analysis [96] have enabled the creation of PDMS microdevices with abrupt and hyperbolic constrictions that mimic in vivo conditions. Shelby et al. [94] pioneered the use of constriction microchannels to measure the deformability of malaria-infected red blood cells (RBCs), revealing reduced deformability when compared to healthy RBCs. This breakthrough has played an important role in the development of similar devices to investigate the flow and deformability of RBCs [97,98], white blood cells (WBCs) [99], and cancer cells [93]. These devices are classified as either structure-induced or fluid-induced deformation channels, depending on the relative size of the microchannel to the tested cells. More details about the flow of cells through structure- and fluid-induced deformation microchannels can be found in the review article performed by Bento et al. [38]. Examples of healthy and pathological cells flowing and undergoing deformation in microfluidic devices with fluid-induced deformation microchannels are shown in Figure 4. It is evident that the RBCs’ deformation is influenced by both the strain and shear rates.
The significant shear effects in microfluidic devices with structure-induced deformation microchannels lead to the substantial deformation of RBCs. Despite their great popularity, the small size of these microchannels may pose a high number of operational problems and challenges. Fluid-induced deformation microchannels could be a valid solution to overcome these issues. This approach has the influence of both shear and extensional flows and is much easier to fabricate. Such microchannels are characterized by abrupt, sudden, or hyperbolic constrictions [38,100,101]. Studies by Zhao et al. [97] and Bento et al. [38] have successfully demonstrated the application of PDMS microfluidic devices with abrupt or sudden microcontractions. Zhao et al. [97] examined RBC deformability at various flow rates in a PDMS microchannel with an abrupt contraction, revealing that RBC elongation reaches a maximal value and then ceases to deform. Lima et al. [102] have developed a cost-effective, user-friendly particulate blood analogue, and they compared the deformability of micelles and RBCs in a sudden-contraction microchannel.
During the last decade, microfluidic devices with hyperbolic constrictions have gain increased popularity as they offer a more stable extensional flow compared to abrupt constrictions [38,101,103]. Note that, by using hyperbolic constrictions, the extensional flow can be more homogeneous as the region where the cells are measured are subjected to a constant strain rate [38,104]. This geometry facilitates not only the study of the RBCs tumbling, rolling, and tank-treading motion, but also the deformability assessment of both healthy [103] and pathological RBCs [83,100]. Figure 4a shows schematically RBC deformation behavior through a hyperbolic microconstriction at two flow rates and distinct locations. As previously noted by Zhao et al. [97] and others [38], RBC elongation in sudden contraction microchannels reaches a maximum value, with no further deformation observed. Conversely, Zeng and Ristenpart [105] found that that RBC deformability tends to decrease slightly as they progress within the contraction region. These contradictory findings highlight the need for further research on the micro-scale blood flow phenomena, such as cell-free layer (CFL), RBC interaction, orientation, and deformability.
3.1.2. Application in Microfluidic Devices with Bifurcations
The ability of PDMS to culture cells on microchannel surfaces is one of its the main advantages [106,107,108,109]. Another remarkable advantage of PDMS is its capacity to produce intricate and complex geometries at both macro- and micro-scale levels. In fact, one of the first works to assess the ability to culture endothelial cells in confined microvascular networks was performed by Shin et al. [110]. This work demonstrated that endothelial cells can be cultured in microvascular bifurcations, which was an important step towards the development of in vitro vasculature for OoC systems. Geometries such as bifurcations and confluences can be found in PDMS microfluidic devices and several blood flow studies have been performed from the beginning of the current century [111] up to now [39,112,113,114]. These studies have shown the importance of understanding how these intricate microvascular networks can affect blood flow behavior and how they can be used to manipulate blood samples for biomedical purposes. Blood cell flow across bifurcation channels is dependent on several factors at the micro-scale level, such as cell size and dispersion at the parent channel [115], hematocrit distribution [116], and cell deformability and aggregation [117].
Blood flow in microfluidic devices typically exhibits unique flow patterns and rheological features, such as the cell-free layer (CFL) on the walls and a high cell concentration in the center region [39,71,118]. However, using PDMS microfluidic platforms produced by xurography [5,39] and soft lithography [39,119] has revealed CFLs at both the walls and confluence apex region. Bento et al. [111] have performed several blood flow studies about the formation and behavior of the CFL in PDMS microfluidic devices with microchannel networks [120]. The results indicate that hematocrit (Hct) has a big impact on the CFL. Additionally, CFLs are most likely to be found on the walls and just downstream of the confluence apex in the middle region of the microchannel networks (see Figure 5a). This study clearly demonstrates that this flow phenomenon occurs in PDMS microfluidic devices with microchannel networks having rectangular cross-sections and with low and moderate Hcts (1 to 15% Hct). Furthermore, Bento et al. conducted another research work on microchannel networks [121] and examined the motion of microbubbles and their impact on local Hct. The findings demonstrate that air bubble passage has a significant impact on the local Hct, as higher cell concentrations were seen upstream of the bubble and lower local Hcts were seen downstream of the bubble (see Figure 5b). In addition, they have found that asymmetric bubble splitting at the bifurcation areas could result in an unequal cell distribution along the outflow branches.
In addition to the experimental blood flow studies performed in PDMS microfluidic devices, it is also fundamental to develop and improve numerical flow models [122,123,124,125] in order to improve our understanding regarding flow behavior in devices with complex geometries such as microchannel networks [114,115] and OoC platforms [126]. Belenkovich et al. [114] have performed both numerical and experimental flow analysis in a PDMS bifurcating microchannel network based on the Murray law, where the wall shear rate remains constant throughout the microchannels. By using these strategies, they have demonstrated that the regions located at the divergent bifurcations of microchannels tend to facilitate the formation of thrombus when compared to straight channels (see Figure 6a). Carvalho et al. [126] have used an OoC device, fabricated in PDMS, and have experimentally validated a numerical model capable of reproducing the fluid flow behavior within an OoC. The numerical model validated in this study showed the potential to accurately predict and evaluate the fluid flow within the OoC device (see Figure 6b). These results demonstrate the importance of validating numerical models. Once validated, the numerical simulations can be a valuable tool to minimize both costs and time of the OoC design by reducing the need to produce prototypes and perform preliminary laboratory experiments.
3.1.3. Application of PDMS Based Blood Analogues in Microfluidics
Blood analogues are fluids usually used in blood flow experiments due to safety concerns with real blood. Early analogues, like glycerol–water or xanthan gum mixtures, could not replicate micro-scale phenomena like CFL, plasma skimming, and cell margination. These effects require particulate blood analogue fluids with solid elements such as microparticles or microcapsules. Recent research has focused on developing particulate blood analogues with varied stiffness, shape, and size for biomedical applications, enabling better replication of microcirculation phenomena [12,21].
PDMS-based blood analogues have gained attention due to their flexibility and unique mechanical properties. Several research works have proposed different kinds of methods to produce flexible PDMS microparticles for biomedical applications, such as flow-focusing techniques, emulsification methods, and multi-stage membrane emulsification processes [12,21]. More details about PDMS-based blood analogues in microfluidics can be found in the comprehensive reviews performed by Sadek et al. [21] and Miranda et al. [12].
3.2. Application of PDMS to Produce In Vitro Biomodels
The first in vitro biomodels (also known as phantoms) were initially produced with materials like glass [127,128], latex [129], and polymethyl methacrylate (PMMA) [130,131], but their rigidity, fabrication complexity, and high costs have reduced their use to fabricate biomodels. Recently, PDMS has gained attention from the biomedical community due to its excellent mechanical properties, and the ability to replicate complex cardiovascular geometries with high resolution [65,132,133,134]. However, manufacturing PDMS biomodels poses challenges, especially in replicating flexible geometries with thin arterial walls. Additive Manufacturing (AM) is a commonly used fabrication process, but it cannot directly print PDMS due to its curing process and high viscosity. Thus, a hybrid process combining AM and PDMS casting is frequently employed [134]. The process to obtain a PDMS biomodel usually involves a 3D printer to fabricate the artery lumen model, then the model is placed in a container or counter-mold, and at the end, the printed material is removed to produce a transparent PDMS biomodel or phantom. Figure 7 shows a schematic overview of this process.
Recent advances in manufacturing biomodels of aneurysms, bifurcations, and stenoses have demonstrated significant progress in material selection, fabrication techniques, and validation methods. These developments have improved our understanding of vascular pathologies and hemodynamic phenomena, facilitating both experimental and computational studies. The process to fabricate intracranial aneurysm (IA) models often combines 3D printing and soft lithography. One study used clinical data to produce a simplified IA model via CAD software and then, by using a FDM 3D printer, molds were produced in Acrylonitrile Butadiene Styrene (ABS). Afterwards, PDMS was poured into the ABS molds, resulting in a flexible IA model. This approach enabled the study of wall deformation under different flow rates using Digital Image Correlation [5,135,136]. In order to obtain transparent models at low cost, Falk et al. [137] developed biomodels using polyvinyl alcohol (PVA) by using a lost-core casting technique. These models were suitable for performing Particle Image Velocimetry (PIV) experiments. Similarly, Souza et al. [134] fabricated IA models using a stereolithography (SLA) printer and a lost-core molding technique with paraffin, beeswax, and glycerin soap. All PDMS biomodels have shown excellent transparency and reproducibility, and they were validated through dimensional analysis and particle visualization tests. More recently, Souza et al. [65], by using a PDMS biomodel, have assessed the impact of hemodynamics on a real intracranial aneurysm (IA), using both experiments and computational fluid dynamics (CFD) simulations. A particle tracking velocimetry (PTV) approach was used not only to study the vortical structures inside the IA, but also to validate numerical simulations performed at a steady regime for different flow rates (see Figure 8). This study has shown great potential for combining PTV and CFD in order to acquire detailed insights into flow structures within IA aneurysms, which are crucial for evaluating the most effective treatments and interventions. Another recent study conducted by Karam et al. [138] employed a digital light processing (DLP) printer with a water-dilutable resin to create patient-specific IA biomodels. Micro-CT scanning and refractive index tests confirmed their geometric accuracy and transparency, respectively, making this a promising method to produce low-cost PDMS biomodels for in vitro blood flow research. Table 3 presents developed biomodels for in vitro blood flow studies.
PIV is widely used to experimentally study in vitro flow dynamics in biomodels. This technique measures velocity fields by tracking trace particle displacement over time with high-speed cameras, providing invaluable insights into hemodynamic behavior and validating computational fluid dynamic (CFD) models. Ford et al. [139] conducted a study comparing CFD-predicted velocity fields with those measured using PIV. Their experiments used anatomically realistic biomodels of a giant aneurysm in the internal carotid artery and another aneurysm at the basilar artery tip, made of transparent PDMS. The study revealed a strong correlation between PIV measurements and CFD predictions, demonstrating PIV’s efficacy in validating numerical models and studying aneurysm geometry-related risk factors. Doutel et al. [140] used a rapid prototyping technique for producing PDMS biomodels using lost-core casting with sucrose. This biomodel provided excellent optical access for performing micro-Particle Image Velocimetry (µPIV) [140]. Employing this method, Doutel et al. investigated geometry’s influence on blood flow in vitro and in silico [141]. Comparative analyses of idealized and patient-specific coronary artery models revealed significant discrepancies. Idealized models displayed uniform narrowing, whereas patient-specific models exhibited uneven narrowing, highlighting the necessity of patient-specific modeling in hemodynamic studies. Jewkes et al. [142] fabricated PDMS 3D models of healthy and stenotic porcine coronary arteries suitable for flow studies. Preliminary flow visualizations revealed helical flow patterns in healthy arteries and recirculation zones in stenotic models, with the latter more prominent in diseased cases. Despite these insights, limitations such as water as a working fluid and using a mobile phone to perform the flow visualizations potentially compromised accuracy. Kefayati et al. [143,144,145] conducted extensive studies on stenotic bifurcation carotid arteries, integrating PIV and CFD analyses. They fabricated PDMS flow phantoms by using lost-core casting. Three artery configurations were examined: a healthy model and two stenotic models with 50% and 70% stenosis, respectively. Initial investigations identified transitional flows using PIV [143]. Subsequent studies explored stenosis severity (30%, 50%, and 70%), plaque eccentricity, and ulceration [144]. The results highlighted increased turbulence with stenosis severity and the pronounced effects of eccentricity and ulceration. Further analysis revealed that shear stress levels tend to increase with stenosis severity and they have found distinct differences between concentric and eccentric plaques [145]. These findings emphasize the clinical implications of parameters beyond stenosis severity in stroke risk assessment. Choi et al. [146] compared rigid and flexible neck artery constrictions using PIV. Their flexible constriction model, fabricated via 3D printing, replicated realistic pulsatile blood flow conditions. To simulate the thin fibrous layer, careful adjustments were performed in PDMS composition. During the assessment of the effect of stenotic deformation on the pulsatile waveform and pressure drop, they observed that flexible constriction tends to change its shape in response to inflow variations. The findings indicate that it is possible to use the pressure drop waveform as a means of identifying susceptible flexible constrictions.
Table 3Summary of some rigid and flexible biomodels for in vitro blood flow studies.
| Geometry | Fabrication Method and Material | Cast | Blood Analogue | Measurement Method | Ref. |
|---|---|---|---|---|---|
| Real intracranial aneurysms | Stereolithography (SLA); photopolymer resin | PDMS | Dimethyl sulfoxide (DMSO) in water | PTV | [65,134] |
| Real aneurysms | FDM 3D printer; ABS | PDMS | Water-Glycerin | PIV | [139] |
| Intracranial aneurysm | Digital light processing (DLP) printer; resin | PDMS | Water–Glycerin–Urea | [138] | |
| Carotid artery | Lost-core manufacturing technique | PDMS | Water–Glycerin–Sodium iodide | Stereoscopic PIV | [143,144,145] |
| Neck artery constriction | 3D printer | PDMS | Water–Glycerin | PIV | [146] |
| Coronary artery | Lost-core casting with sucrose | PDMS | Water–Glycerin | Micro-PIV | [20,140,141] |
| Intracranial aneurysm | FDM 3D printer; ABS | PDMS | Water–Glycerin | Digital Image Correlation | [136] |
| Porcine coronary arteries | 3D printer | PDMS | Water | Mobile phone | [142] |
3.3. Application of PDMS for Heat Transfer Studies
PDMS has been used in the development of several innovative devices and it plays a critical role in enhancing the performance of heat exchangers, microchannel heat sinks, microelectronic systems, pulsating heat pipes, and thermal conductivity measurement devices.
PDMS has been employed to create pulsating heat pipes (PHPs), which are essential for electronics cooling [147]. A study involving the fabrication of miniaturized PDMS-based PHPs showed that PDMS’s flexibility and low-cost manufacturing made it an excellent material for compact heat dissipation devices. Using ethanol and methanol as working fluids, the PHPs demonstrated efficient heat transfer, particularly when oriented vertically. The vertical orientation has improved fluid flow and reduced thermal resistance, especially when methanol was used. The results highlighted the role of PDMS in creating low-cost, efficient cooling systems for electronic devices, particularly 3C products (computers, communication, and consumer electronics) [147].
Souza et al. [56] developed a serpentine-shaped PDMS heat exchanger. This PDMS heat exchanger was made by using 3D-printed molds (see Figure 9a) and was compared with traditional copper-based models. The PDMS serpentine structure maximizes the contact area between the cooling fluid and the heated surface, improving significantly the thermal exchange efficiency in applications such as photovoltaic solar panels, CPUs, and other heat-sensitive electronics. Comparative tests have shown that the PDMS heat exchanger reduces surface temperature by an average of 17% compared to traditional copper-based systems (see Figure 9b). This study has shown that PDMS can be a promising cost-effective solution not only to improve thermal exchange efficiency, but also to perform precise thermal management in electronics, solar panels, and super computers. This work resulted in patent PT118128(A) [148].
Microchannel heat sinks are essential for performing efficient heat dissipation in compact electronic devices. PDMS’s flexibility and simple fabrication have facilitated the creation of microchannel structures that enhance forced convective heat transfer [149]. Jung et al. [149] employed μPIV-μLIF techniques and found that PDMS microchannels have the ability to perform precise measurements of temperatures and velocity fields, enabling the analysis of flow behavior at various Reynolds numbers (Re). This study found that higher Re values led to higher crossflow and spanwise vorticity, resulting in improved heat transfer [149]. The ability to visualize and optimize these flow dynamics in PDMS microchannels will be useful in improving our understanding regarding the flow of thermofluids associated with heat transfer in microchannel heat sinks.
Chuang and Wereley [150], by mixing metallic powders with the PDMS matrix, created a PDMS conductive composite that was used to evaluate its potential for heating and temperature sensing. The integrated microheaters within the PDMS were found to be more feasible as a fixed heating source than running thermal cycle. Regarding the PDMS sensor, the response to temperature changes was slow mainly due to poor thermal conductivity [150]. Nevertheless, they showed reasonable stability at a fixed temperature, making them suitable for some applications with constant heating zones, such as continuous flow PCR systems [151].
Yi et al. [152] developed PDMS nanocomposites by incorporating thermally conductive alumina (Al2O3) nanoparticles. This approach has significantly enhanced the thermal conductivity of PDMS while preserving its flexibility and decreasing the specific heat capacity of the developed materials. Their proof-of-concept study shows that PDMS nanocomposites made of highly thermally conductive nanoparticle offers a promising solution for thermal management and cooling embedded microelectronic systems.
Another interesting application of PDMS is its ability to perform thermal conductivity measurements on nanofluids (NFs). The traditional methods to measure NF thermal conductivity do not allow the assessment of critical features that affect the heat transfer performance of NFs. Aggregation, sedimentation, and NP wall adhesion are some of the crucial features that frequently affect the thermal properties of NFs. Furthermore, because of design, geometry, and material constraints, the traditional cells have critical functional limitations in terms of full cleaning and performing direct visualizations. These are frequent problems encountered at the transient hot-wire and transient plane source (TPS) methods, two popular techniques often used to measure the thermal conductivity of NFs [153,154]. Recently, Souza and his colleagues [153] developed a PDMS cell to improve the thermal conductivity measurements of NFs. Unlike conventional materials, PDMS can reduce nanoparticle adhesion on cell walls, facilitating the cleaning of the cell and consequently demonstrating more consistent results. Hence, the repeatability and consistency of the measurements are significantly improved due to the reduced risk of contamination. In this work, Souza et al. [153] used the developed PDMS cell to measure the NF thermal conductivity through a TPS method (see Figure 9a). Due to its optical transparency and simple handling, this PDMS cell has the unique ability to perform direct visualizations and, consequently, to detect and remove air bubbles and possible NP sedimentation (see Figure 10b). By using this cell, it will be possible to eliminate these potential sources of errors. Unlike traditional cells, this PDMS cell consists of two separable parts (see Figure 10a), ensuring full cleaning and preventing sample contamination. Additionally, the PDMS low thermal conductivity provides natural insulation, minimizing heat exchange with the external environment. These features make this PDMS cell a reliable and efficient alternative method to measure the thermal conductivity of NFs.
Figure 10(a) Procedure to measure the thermal conductivity of nanofluids using fabricated PDMS cell and the Hot Disk TPS 2500S data acquisition system; (b) visualization of air bubbles and sedimentation in a NF with a high concentration (adapted from [153]).
[Figure omitted. See PDF]
3.4. Application of PDMS to Produce Face Masks
The World Health Organization (WHO) in 2020 declared COVID-19 a pandemic, recommending measures such as hand hygiene and the use of face masks to reduce viral spread [155,156]. Face masks have proven essential in reducing transmission rates and to control both public and personal health. In this way, taking into account the progress in microfabrication and biomicrofluidics, a novel transparent face mask was developed using polydimethylsiloxane (PDMS) combined with textile fabrics [157]. A key innovation of the proposed face mask is its sustainable use of PDMS at the end of its life cycle. Used PDMS was ground into small particles (~1 mm) and incorporated into fresh PDMS at a 5:1 prepolymer-to-curing agent ratio. The mixture was poured into an acrylic mold to produce a PDMS transparent window to be placed in the central region of the mask. After curing for one hour in an oven at 80 °C, the resulting PDMS film was sewn onto a three-layer textile fabric. To prevent viral penetration, additional PDMS was applied to seal all the remaining holes. Figure 11 shows a schematic diagram of the different stages to manufacture the proposed face masks.
Face masks aim to protect public health and reduce viral transmission, especially among healthcare workers and the general public. In Europe, face masks must meet the standards of the European Directive EN 14683:2019 [158], including breathability and bacterial filtration efficiency (BFE) tests. Hence, the developed PDMS mask was tested and certified at the Equilibrium laboratory. The breathability test has shown that the mask had respiratory resistance well below the maximum 40 Pa/cm2 threshold, ensuring comfort and airflow for users. The BFE results indicated filtration rates near 90%, below the minimum requirement for level 1 masks used by healthcare professionals. Thus, the proposed PDMS mask qualified as a level 2 mask, suitable for public use in high-contact environments. Another critical feature of the mask is its transparent PDMS window, enabling visibility of the user’s mouth. Transparency tests revealed no significant difference in transmittance (~90%) between pure PDMS and PDMS containing 10% of recycled particles, confirming that recycled PDMS maintains excellent optical transparency. This sustainable PDMS face mask offers a viable, environmentally friendly solution with effective breathability, adequate filtration, and superior transparency for general public use. From this work resulted in patent PT118128(A) [159].
3.5. Additional PDMS Applications in Engineering
While this review primarily focuses on the application of PDMS in biomicrofluidics, heat transfer studies, and face mask fabrication, it is also important to highlight its diverse applications across other engineering fields. PDMS has been employed in tissue engineering due to its biocompatibility and flexibility, which supports the development of scaffolds and molds for cell growth and tissue regeneration [160,161]. In medical implants, PDMS is one of the most successful polymers due to its biocompatibility, excellent resistance to biodegradation, and flexibility [12]. Additionally, PDMS’s selective gas permeability and elasticity make it a preferred material in membrane technologies for gas separation and pervaporation [162,163]. Its role as a coating and sealant offers protection against moisture and chemicals in industrial applications, while its versatility, combined with its excellent mechanical, thermal, and chemical properties makes PDMS an ideal choice for advanced fields like soft robotics and flexible electronics [164,165]. For readers seeking a deeper understanding of these applications, several reviews are available. For example, the reviews conducted by Ariati et al. [163] and Miranda et al. [12] provide insights into the applications of PDMS in medical implants and membrane technologies. Similarly, the reviews of Majidi [164] and Li et al. [165] offer valuable insights into the applications of PDMS in soft robotics and flexible electronics.
4. Conclusions
Polydimethylsiloxane (PDMS) is a versatile silicone elastomer widely used in engineering and biomedical applications due to its unique properties, including optical transparency, gas permeability, biocompatibility, and low-cost fabrication. These characteristics have made PDMS a natural choice for microfluidic devices, organ-on-chip platforms, in vitro biomodels, and particulate blood analogues. PDMS’s flexibility and ability to replicate complex geometries have promoted the progress of the research in areas like disease modeling, drug testing, heat transfer, and hemodynamics.
Despite its benefits, PDMS’s hydrophobic nature poses challenges in biological applications, such as the absorption of hydrophobic molecules and fluid transport in microchannels. Surface treatments like oxygen plasma treatment and surfactant-based modifications have shown promise, but the temporary nature of these treatments remains a limitation. Additionally, the mass production of PDMS-based devices remains a significant drawback, as the current fabrication processes, while suitable for laboratory-scale and small-batch production, face challenges when scaled up to industrial level. These challenges include the recent rising cost of raw materials, time-intensive casting and curing steps, the need for specialized equipment, and, in some cases, the requirement for clean-room facilities. Overcoming these limitations is essential for advancing PDMS and to enable PDMS to reach industrial-scale applications.
Future research should focus on improving hydrophilic treatments to achieve long-term effects and develop scalable manufacturing processes. Exploring hybrid fabrication techniques, including additive manufacturing and nanocomposite integration, could open new avenues for PDMS application. Furthermore, PDMS also shows great potential in several emerging engineering fields like acoustofluidics, in vitro biomodels, and advanced heat transfer systems, highlighting the need for further research. By addressing these challenges and exploring its multifaceted properties, PDMS can continue to revolutionize biomedical engineering and beyond.
I would like to express my gratitude to everyone who directly or indirectly contributed to the success of this research on the applications of PDMS in different fields of engineering.
The author declares that he has no known competing financial Interests.
Footnotes
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![View Image - Figure 1. (a) Most common materials used in microfluidics placed according to the Young’s (elastic) modulus (modified from [3]); (b) temporal evolution of the main polymers used in microfluidic devices, highlighting PDMS and thermoplastics such as PS and PMMA (modified from [11]).](https://media.proquest.com/media/hms/THUM/2/s14jb?_a=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%3D%3D&_s=m5LM7VBlX%2BU%2FwxObYFyrxxysq0I%3D "View Image - Figure 1. (a) Most common materials used in microfluidics placed according to the Young’s (elastic) modulus (modified from [3]); (b) temporal evolution of the main polymers used in microfluidic devices, highlighting PDMS and thermoplastics such as PS and PMMA (modified from [11]).")
Enlarge this image.Figure 1. (a) Most common materials used in microfluidics placed according to the Young’s (elastic) modulus (modified from [3]); (b) temporal evolution of the main polymers used in microfluidic devices, highlighting PDMS and thermoplastics such as PS and PMMA (modified from [11]).
Enlarge this image.Figure 2. PDMS applications in different fields of engineering including biomicrofluidics, blood analogues, organs-on-a-chip, in vitro biomodels, heat transfer, and face masks.
![View Image - Figure 3. Schematic capillary test on a PDMS microfluidic device after 420 h of treatment by O2-plasma-PEG [80]. It is shown that at 8 s, the fluid with Rhodamine B was halfway through the microchannel, and at 13 s, the microchannels were completely filled. With untreated PDMS, no flow was observed.](https://media.proquest.com/media/hms/THUM/8/s14jb?_a=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%3D%3D&_s=237CiZMcI8XItWk0wECnsTaPe%2BM%3D "View Image - Figure 3. Schematic capillary test on a PDMS microfluidic device after 420 h of treatment by O2-plasma-PEG [80]. It is shown that at 8 s, the fluid with Rhodamine B was halfway through the microchannel, and at 13 s, the microchannels were completely filled. With untreated PDMS, no flow was observed.")
Enlarge this image.Figure 3. Schematic capillary test on a PDMS microfluidic device after 420 h of treatment by O2-plasma-PEG [80]. It is shown that at 8 s, the fluid with Rhodamine B was halfway through the microchannel, and at 13 s, the microchannels were completely filled. With untreated PDMS, no flow was observed.
![View Image - Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).](https://media.proquest.com/media/hms/THUM/10/s14jb?_a=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%3D%3D&_s=fFHv7Q6p3cuKuysO4U47h6%2FB7F8%3D "View Image - Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).")
Enlarge this image.Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).
![View Image - Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).](https://media.proquest.com/media/hms/THUM/12/s14jb?_a=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%3D%3D&_s=lTlADPxm90Ja30c1NRMEGhNJCvw%3D "View Image - Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).")
Enlarge this image.Figure 4. RBCs flowing through (a) a microchannel with a hyperbolic constriction; (b) a multi-step microfluidic device able to separate RBCs and WBCs from plasma and simultaneously measure blood cell deformability (modified from [5,92,100]).
![View Image - Figure 5. (a) CFL visualization in a microchannel network made of PDMS (adapted from [120]); (b) air bubble effect on the local Hct for a flow rate of 10 µL/min and 10% Hct (from [121]).](https://media.proquest.com/media/hms/THUM/14/s14jb?_a=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%3D%3D&_s=kvjzbbASq%2Fhn91oYPhg2XxQWt%2BA%3D "View Image - Figure 5. (a) CFL visualization in a microchannel network made of PDMS (adapted from [120]); (b) air bubble effect on the local Hct for a flow rate of 10 µL/min and 10% Hct (from [121]).")
Enlarge this image.Figure 5. (a) CFL visualization in a microchannel network made of PDMS (adapted from [120]); (b) air bubble effect on the local Hct for a flow rate of 10 µL/min and 10% Hct (from [121]).
![View Image - Figure 6. (a) Computational fluid dynamics showing the wall shear rate (WSR) and the velocity field for straight (top) and Murray’s law-based models (bottom), and experimental results in bifurcation microvessel networks (adapted from [114]). (b) Computed 3D velocity profiles in an OoC device, velocity streamlines, and wall shear stress (WSS) around organoids (adapted from [126]).](https://media.proquest.com/media/hms/THUM/16/s14jb?_a=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%3D%3D&_s=s7QuaIk%2BBF2XsX2NW7PM%2B5FQGkg%3D "View Image - Figure 6. (a) Computational fluid dynamics showing the wall shear rate (WSR) and the velocity field for straight (top) and Murray’s law-based models (bottom), and experimental results in bifurcation microvessel networks (adapted from [114]). (b) Computed 3D velocity profiles in an OoC device, velocity streamlines, and wall shear stress (WSS) around organoids (adapted from [126]).")
Enlarge this image.Figure 6. (a) Computational fluid dynamics showing the wall shear rate (WSR) and the velocity field for straight (top) and Murray’s law-based models (bottom), and experimental results in bifurcation microvessel networks (adapted from [114]). (b) Computed 3D velocity profiles in an OoC device, velocity streamlines, and wall shear stress (WSS) around organoids (adapted from [126]).
![View Image - Figure 7. Schematic diagram of the manufacturing process to obtain PDMS biomodels. The artery lumen model is fabricated using a 3D printer, then placed in a container or counter-mold, and at the end, the lumen is removed to produce a transparent PDMS biomodel (adapted from [5]).](https://media.proquest.com/media/hms/THUM/18/s14jb?_a=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%3D%3D&_s=AQb9aj71%2BD%2FXgz1SuRPV88W2tog%3D "View Image - Figure 7. Schematic diagram of the manufacturing process to obtain PDMS biomodels. The artery lumen model is fabricated using a 3D printer, then placed in a container or counter-mold, and at the end, the lumen is removed to produce a transparent PDMS biomodel (adapted from [5]).")
Enlarge this image.Figure 7. Schematic diagram of the manufacturing process to obtain PDMS biomodels. The artery lumen model is fabricated using a 3D printer, then placed in a container or counter-mold, and at the end, the lumen is removed to produce a transparent PDMS biomodel (adapted from [5]).
![View Image - Figure 8. Experimental (PTV) and numerical results in a PDMS-based intracranial aneurysm (IA) biomodel for different Reynolds number (Re). The impact of hemodynamics on a real IA under steady-state flow and for different flow rates [65].](https://media.proquest.com/media/hms/THUM/20/s14jb?_a=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%3D%3D&_s=5%2BkgNBpmutXhMbUnSGltr3CHvLE%3D "View Image - Figure 8. Experimental (PTV) and numerical results in a PDMS-based intracranial aneurysm (IA) biomodel for different Reynolds number (Re). The impact of hemodynamics on a real IA under steady-state flow and for different flow rates [65].")
Enlarge this image.Figure 8. Experimental (PTV) and numerical results in a PDMS-based intracranial aneurysm (IA) biomodel for different Reynolds number (Re). The impact of hemodynamics on a real IA under steady-state flow and for different flow rates [65].
![View Image - Figure 9. Schematic representation (a) of the method to produce the PDMS serpentine heat exchanger; (b) of the experimental set-up (adapted from [56]).](https://media.proquest.com/media/hms/THUM/22/s14jb?_a=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%3D%3D&_s=u1hvc5Buz8sUHa1N4TBAWbxWhn4%3D "View Image - Figure 9. Schematic representation (a) of the method to produce the PDMS serpentine heat exchanger; (b) of the experimental set-up (adapted from [56]).")
Enlarge this image.Figure 9. Schematic representation (a) of the method to produce the PDMS serpentine heat exchanger; (b) of the experimental set-up (adapted from [56]).
Enlarge this image.Figure 11. PDMS mask: (a) schematic representation of the process to produce face masks with recycled PDMS and textile fabrics; (b) real PDMS face mask.
Relevant PDMS properties for fluid applications.
| Property | Value | References |
|---|---|---|
| Optical transparency | 240–1100 (nm) | [ |
| Hydrophobicity—contact angle | ~108 ± 7 (°) | [ |
| Refraction index | 1.4 | [ |
| Thermal conductivity | 0.2–0.27 (W/m∙K) | [ |
| Specific heat | 1.46 (kJ/kg∙K) | [ |
| Electrical conductivity | 4 × 1013 (ohm∙m) | [ |
| Longitudinal wave velocity | 1028.3–1119.1 (m/s) | [ |
| Shear wave velocity | 75–124.3 (m/s) | [ |
| Young’s elastic modulus | ~1–3 (MPa) | [ |
| Poisson ratio | 0.5 | [ |
| Tensile strength | 2.24–6.7 (MPa) | [ |
| Hardness | 41–43 (Shore A) | [ |
| Density | 1029.4–1031.4 (kg/m3) | [ |
| Viscosity | 3.5 (Pa∙s) | [ |
Main advantages and drawbacks of fabricating microfluidic devices, in vitro biomodels and OoC platforms [
| Materials | Main Advantages | Main Disadvantages |
|---|---|---|
| PDMS | Optical transparency, gas permeability, simple and low-cost fabrication, biocompatibility, variable elasticity, and cell culture. | Can absorb hydrophobic molecules, hydrophobic nature, difficult mass production, and its attenuation of acoustic waves. |
| Hydrogel | Low cost, allows diffusion of small molecules, biocompatibility, and cells can be loaded on the surface or to the bulk. | Degradable, weak mechanical strength, and requires freezing or drying for long-term storage. |
| Thermoplastics | Low-cost fabrication, optical transparency, and mass production. | Rigid, thermal degradation and thermal oxidative degradation in the presence of oxygen, and permeability inability. |
| 3D printing resins | Simple and low-cost fabrication, variable mechanical properties, and ability to create complex geometries. | Inadequate optical transparency, low gas permeability, surface roughness, and limited material choices depending on printer technology. |
| Glass | Optical transparency, inert, and excellent roughness. | Rigid, fragile, expensive, and difficult to reproduce complex geometries. |
| Silicon | Ability to create complex geometries at both micro and nano level, thermal stability, and chemical resistance. | High-cost fabrication, need for clean-room facilities, permeability inability, and no optical transparency. |
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1 MEtRICs, Mechanical Engineering Department, University of Minho, Campus de Azurém, 4800-058 Guimarães, Portugal;