Multimaterial 3D printing hydrogel with other polymers
As shown in Fig. 1A, we print the hydrogel-polymer heterogeneous structures on a self-built high-resolution, high-efficiency DLP-based multimaterial 3D printer (24, 25). The system adopts “bottom-up” projection approach where digitalized UV light is irradiated from the UV projector, which is placed below the printing stage that moves vertically to control the layer thickness of each layer printing. Between the printing stage and UV projector, there is a polytetrafluoroethylene-coated glass plate that supports two to three polymer precursor solution puddles and moves horizontally to deliver a needed polymer precursor solution for the corresponding layer. The hydrogel used here is the highly stretchable and high–water content UV curable acrylamide-poly(ethylene glycol) diacrylate (PEGDA) (AP) hydrogel (27). The UV curable polymers are commercially available (meth)acrylate-based 3D printing polymers. Figure 1 (B and C) demonstrates the process of printing a diagonally symmetric Kelvin form where the half part is made of an elastomer (transparent) and the other half is made of AP hydrogel (yellow). When a layer of the elastomer part is being printed (Fig. 1B), the transparent elastomer precursor solution is placed below the printing stage, and the corresponding UV pattern is illuminated into the transparent elastomer precursor solution. As shown in Fig. 1C, when a layer of the hydrogel part is being printed, the yellow hydrogel precursor solution is positioned below the printing stage, and the corresponding UV pattern is irradiated into the hydrogel precursor solution. The air-jetting approach is applied to remove the precursor solution residuals on the printed part and minimize the material contamination during material exchange. Details about the DLP-based multimaterial 3D printing system can be found in note S1, fig. S1, and movie S1. As demonstrated in Fig. 1D, a diagonally symmetric Kelvin is finally printed by alternatively switching elastomer and hydrogel precursor solutions for each layer. As the strong covalent bonds are formed at the interfaces between the highly deformable AP hydrogel and elastomer, we can compress the printed multimaterial Kelvin form by 50% without finding any debonding between the two constituent materials (Fig. 1E and movie S2). To further demonstrate the versatility of this multimaterial 3D printing approach, we print a Kelvin foam where the three unit cells are printed with rigid polymer, AP hydrogel, and elastomer, respectively (Fig. 1F). Again, the hydrogel forms strong covalent bonds with the other two polymers, and we can stretch the hybrid Kelvin foam structure by large deformation without finding debonding between the unit cells printed with different materials (Fig. 1G and movie S3).
Materials and bonding mechanism
The illustrations in Fig. 2 further introduce the chemicals used to prepare this UV curable, highly stretchable, and high–water content AP hydrogel precursor solution and the mechanism that explains how the printed AP hydrogel could firmly bond with other (meth)acrylate-based UV curable polymers. As shown in Fig. 2A, we prepare the AP hydrogel precursor solution by mixing the acrylamide powders, PEGDA, and the self-prepared water-soluble photoinitiator 2,4,6-trimethylbenzoyl diphenylphosphine oxide (TPO) into water. Here, the PEGDA/acrylamide mixing ratio is 0.625 to 1.25 wt %, and the water content of the AP hydrogel is 70 to 80 wt % (details can be found in Materials and Methods). The mechanical behavior can be tuned by changing the PEGDA/acrylamide mixing ratio, the molecular weight of PEGDA, and the water content (27). Using water-soluble TPO as photoinitiator is the key to making the AP hydrogel precursor solution highly UV curable and 3D printable. To make the hydrophobic commercial TPO powders water soluble, we use the surfactant SDS that contains a hydrophilic head and a hydrophobic tail (Fig. 2B) to encapsulate TPO particles (27, 28). In addition, polyvinylpyrrolidone is used as crystallization inhibitor. After freeze-drying the prepared oil-in-water microemulsions containing TPO, the highly water-soluble TPO nanoparticles are produced. Note that previous works report the approaches of using the commercial photoinitiator Irgacure 819 to prepare UV curable hydrogel precursor solutions (29–32). However, the water insolubility of Irgacure 819 requires users to dissolve it into organic solvent such as ethanol and dimethyl sulfoxide, which is only applicable to the hydrogels with relatively low stretchability and water content (29–32). To 3D print a hybrid structure consisting of AP hydrogel and other polymers, we choose a number of commercially available polymer precursor solutions (details can be found in Materials and Methods). Although the chemical details of these commercial polymer solutions are not given, they are (meth)acrylate-based monomers, cross-linkers, and oligomers with non–water-soluble photoinitiators (Fig. 2C).
Figure 2 (D to J) presents the multimaterial 3D printing process, the resulting chemical structures, and the proposed interface bonding mechanism between AP hydrogel and (meth)acrylate polymer. From Fig. 2 (D and E), the hydrogel precursor solution is solidified to a layer of hydrogel network structure upon a patterned UV projection. The photopolymerization cross-links acrylamide and PEGDA to form a chemically cross-linked network structure (Fig. 2H). Note that there is a small amount of unreacted acrylamide and PEGDA remaining in the polymerized hydrogel structure (Fig. 2E). From Fig. 2 (F and G), we further print a (meth)acrylate polymer layer onto the AP hydrogel layer by projecting a UV pattern into the (meth)acrylate polymer precursor solution, which leads to free radical photopolymerization (Fig. 2I). At the interface, the radicals within the (meth)acrylate polymer can also attack those unreacted monomers and oligomers within the AP hydrogel domain, resulting in the chemical bonding between the (meth)acrylate-based polymer layer and AP hydrogel layer (Fig. 2J and fig. S2).
To validate the mechanism that we propose to explain the interfacial bonding between AP hydrogel and (meth)acrylate polymer, we first carry out Fourier transform infrared (FTIR) spectroscopy to investigate the conversions of AP hydrogel polymerization initiated by different initiators. To compare conversion and kinetics of polymerization, we use another two initiators for comparison: (i) Irgacure 2959 (I2959), a commercial water-soluble photoinitiator that has been widely used in DIW-based acrylamide hydrogel 3D printing (18), and (ii) ammonium persulphate (APS) as thermal initiator plus N,N,N′,N′-tetramethylethylenediamine (TEMED) as initiation accelerator, which have been widely used to synthesize bulk AP hydrogel samples (33). Figure 3A presents the FTIR spectrums of AP hydrogel before and after polymerization. The conversion can be calculated from the decay of the peak at 988 cm−1, which is the out-of-plane bending mode of the ─CH2─ unit (methylene group), normalized to the C═O stretching peak (carbonyl group) at 1655 cm−1 as an internal standard. The conversions are ~100, ~90, and ~10% for the AP hydrogels initiated by APS-TEMED, TPO, and I2959, respectively. Details about sample preparation and FTIR tests can be found in Materials and Methods.
Since both TPO nanoparticle and I2959 are water-soluble photoinitiators, we further perform FTIR to investigate the photopolymerization kinetics of the AP hydrogel initiated by TPO and I2959, respectively. Details about the FTIR tests can be found in Materials and Methods and fig. S3. As shown in Fig. 3B, under the same UV curing condition, the AP hydrogel precursor solution initiated by TPO reaches 80% of conversion in 15 s, and its conversion is saturated at 90% after curing for 2 min. In comparison, the conversion of the AP hydrogel solution initiated by I2959 only reaches 10% after 2 min of UV curing. This is due to the extremely low photosensitivity of I2959 at 405 nm (molar extinction coefficient ε ~2 M−1 cm−1), which is about 300 times lower than that of TPO (ε ~600 M−1 cm−1) (28). On the basis of this, we conclude that AP hydrogel precursor solution with TPO is applicable to DLP-based 3D printing (the wavelength is usually 385 or 405 nm), which requires short UV exposure time (within 1 min) to cure a layer. AP hydrogel precursor solution with photoinitiator I2959 is only suitable to the DIW-based 3D printing, which allows users to print out 3D structures first and conduct a few hours of post-curing under even shorter wavelength UV irradiation to cross-link the hydrogel (18).
From Fig. 3A, we also find that APS-TEMED initiator converts almost 100% of acrylamide and PEGDA to the hydrogel covalent network; in comparison, the conversion of AP hydrogel with TPO is 90%. The difference in the conversions of polymerization is also reflected in the macroscopic mechanical behavior of the cross-linked hydrogels. Figure 3C presents the stress-stretch behavior of the AP hydrogel initiated by APS-TEMED and TPO, respectively. Details about the uniaxial tensile tests can be found in Materials and Methods. The one initiated by TPO exhibits lower stiffness and stretchability as the lower conversion of polymerization results in less cross-linking points and shorter linear chains. Although this relatively lower conversion of polymerization weakens the mechanical performance of the AP hydrogel, it is favorable for bonding the hydrogel with other UV curable (meth)acrylate polymers. To demonstrate that the TPO-initiated AP hydrogel with 90% polymerization conversion can form strong interfacial bonds with (meth)acrylate polymer, which the APS-TEMED–initiated AP hydrogel with 100% polymerization conversion cannot realize, we print a UV curable elastomer (34) onto a TPO-initiated AP hydrogel dogbone sample (Fig. 3D). Since the hydrogel and elastomer form strong chemical bonds, the hydrogel-elastomer hybrid sample can be stretched by five times without debonding. In contrast, the printed UV curable elastomer cannot form strong bonding with the APS-TEMED–initiated hydrogel, and the debonding occurs during stretching the hybrid sample (Fig. 3E). Taking this advantage of the TPO-initiated hydrogel, we further print the TPO-initiated AP hydrogel with other UV curable polymers. As shown in Fig. 3 (F to K), we stretch the hybrid dogbone samples where the AP hydrogel is arranged in series with Tango elastomer, Vero rigid polymer, Agilus elastomer, PEGDA, methacrylate SMP (35), and ABS-like polymer, respectively. Details about the UV curable polymers can be found in Materials and Methods, and uniaxial tensile testing results are presented in fig. S4. Since all the other UV curable polymers are stiffer than the AP hydrogel, most deformations appear on the hydrogel part. The energy needed to break the interfacial bonding between hydrogel and polymer is greater than the energy needed to break the hydrogel itself so that each hybrid sample fails when the hydrogel fractures rather than the debonding happens at the interface (Fig. 3, I to Q).
As illustrated in fig. S5A, we perform the 180° peeling tests to investigate the interfacial toughness between hydrogel and UV curable polymer (fig. S5B and movie S4). Notably, the force/width-displacement curves in fig. S5C indicate that the interfacial toughnesses for the AP hydrogel on the six different UV curable polymers in Fig. 3 (F to K) are similar to each other. This similarity can be explained by the images of the six fractured hydrogel-polymer hybrid samples during the peeling tests (fig. S5D). It can be seen that the AP hydrogel, instead of the hydrogel-polymer interface, undergoes a cohesive failure near the interface during the peeling tests, which leaves a residual layer of hydrogel on the polymer substrates (fig. S5D). The peeling tests confirm that the energy needed to break the hydrogel-polymer interface is greater the energy needed to break the hydrogel itself.
To further investigate the effect of the key printing parameter–UV curing time on the interfacial toughness, we also perform the peeling tests on the AP hydrogel–Tango elastomer hybrid samples where the AP hydrogels are cured with different time. As shown in fig. S6 (A and B), the interfacial toughness gradually increases from 11.3 to 19.7 J/m2 (N/m) by increasing the UV curing from 10 s to 1 min, and it levels off when the UV exposure is longer than 1 min. The increase in the measured interfacial toughness is mainly due the increase in the fracture toughness of the pure AP hydrogel sample by increasing the UV curing from 10 s to 1 min (fig. S6C) as more monomers and cross-linkers are cross-linked. However, the variation in UV curing time does not affect the facture mode. As shown in fig. S6 (D to I), all the hydrogel-elastomer hybrid samples facture on the hydrogel side, and all fractures are cohesive even for the hybrid sample where the hydrogel is only cured for 10 s. In addition, we also perform the peeling test to investigate the interfacial bonding between the Tango elastomer and AP hydrogel initiated by photoinitiator I2959, which is cured in an oxygen-free environment under 254-nm-wavelength UV curing for 40 min. Figure S7 shows that the I2959-initiated AP hydrogel also undergoes cohesive failure, indicating that the proposed interfacial bonding mechanism also works for the I2959-initiated AP hydrogel, which could be used to print AP hydrogel with other polymers through DIW-based 3D printing.
Printed SMP stent with drug releasing function
As illustrated in Fig. 5A, SMP is an ideal 4D printing material (32) for 3D printing shape memorable cardiovascular stent to expand the blood vessel with stenosis. By taking advantage of multimaterial 3D printing, we can impart the drug releasing function into a cardiovascular SMP stent by integrating hydrogel into the SMP stent. Figure 5B presents the design of the multimaterial stent where the major stent structure is printed with SMP, and hydrogel skins (thickness: 200 μm) are embedded into the SMP rods (Fig. 5C). As illustrated in Fig. 5D, since the stent has the shape memory effect, it can be squeezed into a compacted shape at the programming temperature, and the compacted shape can be fixed after cooling to a lower temperature, which is lower than the glass transition temperature (Tg) of the SMP; the compacted stent recovers to its original shape after heating back the programming temperature. As shown in Fig. 5E, we modify the commercial UV polymer VeroClear to reduce its Tg to 30°C (for details, see Materials and Methods) so that we can program the SMP-hydrogel stent at 37°C and the compacted shape can be fixed at 20°C (the shape memory cyclic test is presented in fig. S9). As presented in Fig. 5F, we use the self-built multimaterial DLP printer to fabricate the SMP-hydrogel stent. We load red dye as “drugs” into the AP hydrogel precursor solution (red dye concentration: 2.5 mM/liter). The stent can be fixed to the compacted shape due to the shape memory effect (Fig. 5G). Moreover, as shown in Fig. 5H, the fabrication flexibility of multimaterial 3D printing allows us to vary the size of the SMP-hydrogel stents to adapt the size variation of different blood vessels.
Figure 5I presents an example to demonstrate both shape memory and drug releasing functions of the printed SMP-hydrogel stent. We fabricate a polydimethylsiloxane (PDMS) tube whose diameter varies to a smaller value in the middle to simulate the vessel stenosis. The tube is filled with 37°C phosphate-buffered saline (PBS) solution to mimic the body environment. The expansion of the stent is observed 2 min after it is inserted into the PDMS tube (movie S5), and the stent fully expands after 1 hour, which is consistent with the shape memory cyclic test in fig. S9. After we insert the stent to the tube, the hydrogel starts to release the drugs into the aqueous environment, and the drug releasing process is observed by seeing the color change of the PBS solution. We further quantify the drug releasing process through UV-visible spectrum (note S10 and fig. S10). In Fig. 5J, the hydrogel releases 3, 16, and 30% of the total drugs within 2 min, 30 min, and 1 hour, respectively. After 3 hours, the cumulative release saturates at about 90%. As shown in fig. S11, after being placed in the 37°C PBS solution for 24 hours, the SMP-hydrogel SMP releases the drugs completely, and the SMP rod can be clearly seen through transparent swollen hydrogel skin. On the basis of the real application requirements, the rate of drug release can be further controlled by tuning the size ratio between the hydrogel mesh and drug particles (42), adopting the hydrogel-particle dual-release mechanism (43), using environmental responsive hydrogels that release drug in response to the change in pH or temperature (44).
Printed flexible electronic devices with ionic conductive hydrogel
The ionic conductivity is another promising property, which allows hydrogels to be used in the applications of flexible electronics (5, 30). However, the traditional manufacturing greatly limits the geometrical complexity of hydrogel structures so that the reliability and functionality of manufactured flexible electronics. The capability of 3D printing hydrogel with elastomer offers a promising solution to address this challenge. As shown in fig. S12, we print a 3D ionic conductive hydrogel lattice structure with elastomer protective skin, which protects the structure from water evaporation. Moreover, we can use the capability of 3D printing hydrogel with elastomer to fabricate more complex conductive hydrogel–integrated devices and machines. To demonstrate this, we print a soft pneumatic actuator (SPA) where a conductive hydrogel–based strain sensor is integrated. Figure 6A presents the schematic design of the actuator, and Fig. 6B shows the printed SPA with the hydrogel sensing circuit highlighted in pink color. In comparison with previously reported approaches (see table S2) (45–48), we only need one single step to fabricate the SPA with strain sensor due to the strong interfacial bonding between the UV curable hydrogel and elastomer. To achieve the relation between resistance and deformation, we print the hydrogel sensing circuit on an elastomer substrate. We stretch the substrate along the length direction to characterize the relation between resistance and tension and stretch the substrate along the width direction to capture the relation between resistance and compression (Fig. 6C). We further derive an equation to describe the relation between resistance and stretch: R = aλ2 + bλ−1 with a + b = R0 (initial resistance R0 = 12,249 ohms and fitting parameter a = 11,250 ohms). Details about the derivation can be found in note S9 and fig. S13. Figure 6C shows that the derived equation is able to nicely capture the nonlinear resistance-deformation behavior. Using the derived equation, we can transform the measured resistance variation to the strain on the SPA under bending. As shown in Fig. 6D, when a negative pressure is applied, the bending angle is defined as negative, and the measured strain is positive since the layer where the hydrogel sensor locates is being stretched. The measured strain becomes negative when we apply a positive pressure, which leads to positive bending of the SPA where the hydrogel sensor is being compressed. Note that compared with the previously reported SPA with strain sensor (see table S2) (45–48), we are the first that can measure the compressive strain on the SPA so that the strain sensor developed here can measure the actuations with both positive and negative bending. We further conduct finite element analysis (FEA) (see details in Materials and Methods) to simulate the bending of the SPA and achieve the relation between the bending angle and the strain on the position where the sensor is placed. As shown in Fig. 6 (E and F), the FEA simulations agree with the experiments well, and the bending angle and the strain follow a liner relation. On the basis of this relation between bending angle and strain and the relation between resistance variation and strain, the SPA can sense its instantaneous bending during actuation (49, 50).