Cell-based therapies have garnered significant interest to treat cancer and other diseases. Acoustofluidic technologies are in development to improve cell therapy manufacturing by facilitating rapid molecular delivery across the plasma membrane via ultrasound and microbubbles (MBs). In this study, a three-dimensional (3D) printed acoustofluidic device was used to deliver a fluorescent molecule, calcein, to human T cells. Intracellular delivery of calcein was assessed after varying parameters such as MB face charge, MB concentration, flow channel geometry, ultrasound pressure, and delivery time point after ultrasound treatment. MBs with a cationic surface charge caused statistically significant increases in calcein delivery during acoustofluidic treatment compared to MBs with a neutral surface charge (p < 0.001). Calcein delivery was significantly higher with a concentric spiral channel geometry compared to a rectilinear channel geometry (p < 0.001). Additionally, calcein delivery was significantly enhanced at increased ultrasound pressures of 5.1 MPa compared to lower ultrasound pressures between 0–3.8 MPa (p < 0.001). These results demonstrate that a 3D-printed acoustofluidic device can significantly enhance intracellular delivery of biomolecules to T cells, which may be a viable approach to advance cell-based therapies.

Acoustofluidics (i.e., coupling of ultrasound and fluidic channels) has increasingly been investigated to address challenges in manufacturing and clinical applications, especially in diagnostics and therapeutics. Acoustofluidics technologies have the potential to offer point-of-care (POC) devices for patient bedside use due to improved simplicity and potential for rapid processing. Acoustofluidic research has investigated many potential applications including acoustophoresis (Petersson et al., 2007; Shi et al., 2009; Wu et al., 2018), biomarker analysis (Wang et al., 2020; Zhou et al., 2021), cell concentrator applications (Li et al., 2007; Chen et al., 2014; Kurashina et al., 2017), fluid mixing applications (Shilton et al., 2008; Friend and Yeo, 2011; Yeo et al., 2011), and recently, biomolecular delivery (Belling et al., 2020; Centner et al., 2020). However, the key acoustofluidic parameters that affect biomolecular delivery to cells have not been thoroughly investigated. In addition, the effect of ultrasound-driven microbubble (MB) oscillation on molecular delivery to cells within acoustofluidic channels has not been fully characterized. Optimization of acoustofluidic parameters for biomolecular delivery to cells could potentially enable the use of this technology for non-viral manufacturing of cell therapies and other applications.

Cell-based therapies represent the latest biotechnological revolution in medicine. Potential therapeutic applications include treating cancers, treating autoimmune diseases, treating infectious diseases, and repairing damaged tissue (Buzhor et al., 2014). Recent breakthroughs in adoptive cell therapies, specifically chimeric antigen receptor (CAR)-T cells, have had a significant impact on the treatment of B-cell malignancies, such as leukemia or lymphoma (Buzhor et al., 2014; Kochenderfer et al., 2017; Turtle et al., 2017; Maude et al., 2018), with over 80% of patients achieving complete remission in some trials. CAR-T may also have the potential to treat other forms of cancer and other diseases, such as human immunodeficiency virus (Deeks et al., 2002). CAR-T has indicated the ability to enhance targeted cytotoxic activity while reducing off-target effects due to its specificity, which is a distinct advantage over traditional treatment regimens (e.g., chemotherapy, radiation) for cancer (Moghimi et al., 2021). However, the current manufacturing techniques for CAR-T, which are heavily dependent on viral vectors, have several limitations that limit more widespread adoption of this biotechnology.

Viral vectors, which are infectious entities that are adapted to deliver nucleic acids to the cytoplasm or nucleus, are currently the primary technique utilized for ex vivo modification of T cells (Wang and Riviere, 2015). Viral vectors have been shown to have high efficacy in modifying T cells for clinical trials of new immunotherapies (Ghani et al., 2009; Kochenderfer et al., 2015; Schuster et al., 2017). However, there are several limitations with this technique, including potential insertional mutagenesis and undesired multiplicity of infection (Piscopo et al., 2018). In addition, viral vectors generally cannot deliver non-nucleic acid biomolecules to cells.

To address these limitations, non-viral techniques are being developed that can enhance intracellular molecular delivery. Techniques include nanocarriers and nanoparticles where therapeutic agents may be loading internally or liganded to the surface to improve molecular delivery to specific cells in vivo or in vitro (Majumder et al., 2019). These techniques, however, rely heavily on endocytosis pathways, which may limit molecular delivery efficiency or loading rate by biomodulatory signaling molecules such as mTOR, proton pumps, and cathepsins (Sahay et al., 2013). Additionally, molecular delivery can be reduced via exocytosis of these particles and associated therapeutic agents (Sahay et al., 2013; Oh and Park, 2014). These could have important implications on treatment efficacy as low delivery efficiency or slow loading rates could delay essential cell-based therapies for patients. Another treatment technique includes nanostraw membrane stamping where the plasma membrane is mechanical penetrated. This technique has shown over 85% delivery efficiency of small molecules in adherent cell lines (Zhang et al., 2019). For anchorage independent cells, this technique may be insufficient and has yet to demonstrate feasibility.

Transient permeabilization of the plasma membrane has generated interest since it may allow molecules to enter the cytosol rapidly. Electroporation is a technique that induces transient permeabilization of the plasma membrane and has generated significant interest as a potential method to overcome limitations of viral vectors. However, bulk electroporation may not have the capacity to achieve sufficient levels of molecular delivery that are required for functional applications, such as CAR-T (Annesley et al., 2018). In an attempt to improve delivery efficiency, stimuli-responsive nanoparticles, conjugated nanoparticles, and nanocarriers have been studied in conjugation with electroporation (Kim and Lee, 2017). Another fundamental limitation of electroporation is the risk of epigenetic modifications that can induce a significant increase in cytokine release, which can cause severe immunomodulatory effects after treatment (Falk et al., 2017; He et al., 2019). For example, unintended amplification of critical pro-inflammatory interleukins, such as interleukin-2, may induce cytokine release syndrome, where the immune system inadvertently attacks native tissues and organs, which can be fatal in severe cases. An alternative approach that has been explored for non-viral molecular delivery involves microfluidic squeezing, where cells are passed through microfluidic channels with small constrictions that transiently increase membrane permeability for intracellular delivery of biomolecules (Sharei et al., 2013). However, channel blockage and limited throughput are challenges that can occur with constricted channels.

Acoustofluidics potentially offers an alternative method of biomolecular delivery to T cells, which could address limitations of viral transduction and other approaches. Recent advancements of targeted endonucleases that modify the genome, such as clustered regulatory interspaced palindromic repeats (CRISPR)-Cas9, offers efficient targeting of gene(s) for gene disruption, gene correction, or gene insertion (Morgan et al., 2017). These target genome-editing techniques require alternative intracellular delivery methods as viral vectors have size limitations, which has hindered the encapsulation of targeted endonucleases into a singular viral delivery vehicle (Hirsch et al., 2016; Yang et al., 2016). Acoustofluidic approaches can rapidly deliver biomolecules into cells by transiently increasing membrane permeability (Belling et al., 2020; Centner et al., 2020). Several previous studies have evaluated acoustofluidic-enhanced molecular delivery without ultrasound contrast agents by utilizing acoustic radiation force and shear stress to induce transient perforation in the plasma membrane (Carugo et al., 2011; Belling et al., 2020). In this study, we have characterized acoustofluidic-enhanced molecular delivery with MBs in order to understand the key parameters that affect the efficiency of molecular delivery to human T cells.

The objective of this study was to assess the role of important acoustofluidic parameters on molecular delivery and cell viability. The concepts of sonoporation and ultrasound-mediated molecular delivery to cells are well-established in static systems and have demonstrated significant molecular delivery with consistent treatment parameters (Miller et al., 1999). Recently, a polydimethylsiloxane (PDMS)-based acoustofluidic device was utilized to assess physical effects of MBs on enhancing molecular delivery to in situ tissue (Pereno et al., 2018); however, this approach has not been thoroughly characterized in an acoustofluidic system as a method to enhance molecular delivery in ex vivo applications with lipid-coated MBs. This study aims to provide new insights into the effect of MB phospholipid surface charge, acoustofluidic channel geometry, molecular exposure time point, acoustic pressure, and MB concentration on molecular delivery efficiency. These parameters may have a key role in the development of effective acoustofluidic approaches for manufacturing of cell-based therapies or other molecular delivery applications.

Acoustofluidic devices were fabricated in Accura 60 plastic using stereolithography three-dimensional (3D) printing (Xometry, Gaithersburg, MD). Stainless steel barbed tube fittings (McMaster-Carr, Elmhurst, IL) were inserted into pre-threaded inlet and outlet ports for connection with PVC tubing (1/16 in. ID, 10–32 threads). Three different acoustofluidic channel geometries were designed in SolidWorks (Waltham, MA) and fabricated for comparison of molecular delivery efficiencies [Fig. 1(A)]. Two acoustofluidic designs had rectilinear channels, with square cross-sectional diameters of 1 mm × 1 mm or 2 mm × 2 mm. Additionally, a concentric spiral channel design was also fabricated with a square cross-sectional diameter of 1 mm × 1 mm to increase residence time within the ultrasound beam compared to the rectilinear channel geometry. One millimeter and 2 mm will allow approximately 1.67 wavelengths and 3.33 wavelengths of ultrasound signal in the acoustofluidic channel with a 2.5 MHz center frequency, respectively. Additionally, the selected center frequency (2.5 MHz) is near the MB resonance frequency and should provide sufficient oscillation power with lipid-encapsulated MBs (Versluis et al., 2020).

FIG. 1.

(Color online) (A) Left: Concentric spiral acoustofluidic device design with 1 mm channel width. Middle: Rectilinear acoustofluidic device design with 1 mm channel width. Right: Rectilinear acoustofluidic device design with 2 mm channel width. (B) Experimental configuration of ultrasound transducer (P4-1) placed on top of acoustofluidic device with ultrasound gel for acoustic coupling with the acoustofluidic device. (C) Ultrasound waveform generated by P4-1 transducer as measured by needle hydrophone.

FIG. 1.

(Color online) (A) Left: Concentric spiral acoustofluidic device design with 1 mm channel width. Middle: Rectilinear acoustofluidic device design with 1 mm channel width. Right: Rectilinear acoustofluidic device design with 2 mm channel width. (B) Experimental configuration of ultrasound transducer (P4-1) placed on top of acoustofluidic device with ultrasound gel for acoustic coupling with the acoustofluidic device. (C) Ultrasound waveform generated by P4-1 transducer as measured by needle hydrophone.

Close modal

All experiments were performed with the acoustofluidic devices connected to a peristaltic pump at a flow rate of 1.5 ml/min. The ultrasound transducer was placed directly on top of the acoustofluidic device with ultrasound gel for acoustic coupling [Fig. 1(B)]. A P4–1 ultrasound transducer (96 elements, 2.5 MHz center frequency, ATL, Bothell, WA) was used to transmit B-mode pulses generated by an ultrasound imaging system (Vantage 64LE, Verasonics, Kirkland, WA). The free-field ultrasound pressure output was measured using a calibrated 0.2-mm needle hydrophone (Precision Acoustics, Dorset, UK) 40 mm away from P4–1 ultrasound transducer in a water tank with water as the medium. The –3 dB beam width was 8 mm and the peak negative ultrasound pressure output was 3.8 MPa for all experiments unless otherwise specified [waveform shown in Fig. 1(C)].

Phospholipid-coated gas-filled MBs were custom synthesized to enhance biomolecular delivery to T cells in the acoustofluidic devices. MBs with a cationic surface charge were synthesized using chloroform solutions of 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC, Avanti Lipids, Alabaster, AL); 1,2-distearoyl-sn-glycero-3-ethylphosphocholine (DSEPC, Avanti Lipids); 1,2-distearoyl-sn-glycero-3-phosphoglycerol (DSPG, Avanti Lipids); and polyethylene glycol-40 stearate (Sigma-Aldrich, St. Louis, MO) at a molar ratio of 100:43:1:4.5. The cationic lipid DSEPC conferred a net positive surface charge on the MBs as previously described (Kopechek et al., 2015; Kopechek et al., 2016; Kopechek et al., 2019). MBs with neutral surface charge were synthesized with chloroform solutions of DSPC and 1,2 distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino(polyethylene glycol) –2000] (DSPE-PEG2000, Avanti Lipids) at a molar ratio of 96:4 (Myrset et al., 2011). Chloroform was evaporated under vacuum and an aqueous micellar lipid solution was prepared by adding 1× phosphate buffered solution (PBS) and sonicating (125 W, Qsonica, Newtown, CT) to disperse the lipids. The resulting 10 mg/mL lipid solution was diluted 1:4 in PBS and sealed in a glass vial. The vial head space was filled with decafluorobutane gas (FlouroMed, Round Rock, TX), followed by amalgamation for 45 s at 4350 cpm (DB-338, COXO, Foshan City, China) to form perfluorobutane gas-filled MBs (MBs). This process yields MBs with a mean diameter of 2 μm as previously described (Kopechek et al., 2015).

Acoustic attenuation measurements of MB solutions were conducted to assess effects of MB properties, acoustofluidic channel geometries, and ultrasound parameters on MB destruction. A custom 3D-printed acoustic chamber was designed in SolidWorks [Fig. 2(A)] and fabricated using fused filament deposition with 1.75-mm polylactic acid filament (PLA, McMaster-Carr) with an Ender-3 3D printer (Creality, Shenzhen, China). Two lead zirconate titanate (PZT) transducers (3.3 MHz center frequency, 25 mm diameter, StemInc, Davenport, FL) were aligned opposite each other 31 mm apart and used as a source and receiver to measure acoustic attenuation. MB solutions (2.5% v/v in 3 ml PBS) were added to 4.5-mL disposable polystyrene cuvettes (VWR, Radnor, PA), which were placed in the acoustic chamber at the center of the ultrasound beam [Fig. 2(B)]. A waveform generator (DG822, Rigol, Suzhou, China) was used to generate a 3.3-MHz pulse with 500 cycles at a pulse interval of 10 ms. Received ultrasound signals (300 traces per sample) were acquired with a digital oscilloscope (DS1202Z-E, Rigol) at a sampling rate of 10 MHz and transferred to a desktop computer for processing with matlab (Mathworks, Natick, MA). Representative waveforms are shown in Fig. 2(C). The attenuation coefficient was calculated as previously described (Kopechek et al., 2011). Power spectra were computed for each waveform and averaged to obtain the average power spectrum for each sample. The peak magnitude was determined at 3.30 MHz for each sample, Ss, and for each corresponding reference measurement without MB solution, Sr. The acoustic path length through the sample (inside the cuvette) was 1 cm. Therefore, the acoustic attenuation coefficient in decibels per unit length (dB/cm) can be computed as follows:

αdB=10log10SrSs.
(1)
FIG. 2.

(Color online) (A) Schematic of acoustic attenuation chamber with a recessed 1.3 cm × 1.3 cm cuvette holder to ensure consistent cuvette sample placement. (B) Photo of acoustic attenuation chamber with polystyrene cuvette placed at the center of the ultrasound beam. The acoustic attenuation chamber contains two co-axially aligned single-element transducers, with one transducer acting as the source and the other transducer acting as the receiver. (C) Representative time-domain waveforms acquired using the acoustic attenuation chamber with microbubbles (2.5% v/v) and without microbubbles (reference). Acquired time-domain waveforms were transformed to the frequency domain to calculate acoustic attenuation coefficients using Eq. (1).

FIG. 2.

(Color online) (A) Schematic of acoustic attenuation chamber with a recessed 1.3 cm × 1.3 cm cuvette holder to ensure consistent cuvette sample placement. (B) Photo of acoustic attenuation chamber with polystyrene cuvette placed at the center of the ultrasound beam. The acoustic attenuation chamber contains two co-axially aligned single-element transducers, with one transducer acting as the source and the other transducer acting as the receiver. (C) Representative time-domain waveforms acquired using the acoustic attenuation chamber with microbubbles (2.5% v/v) and without microbubbles (reference). Acquired time-domain waveforms were transformed to the frequency domain to calculate acoustic attenuation coefficients using Eq. (1).

Close modal

Jurkat T cells were cultured with complete RPMI-1640 (10% fetal bovine serum, 1% penicillin/streptomycin) at 37 °C and 5% CO2 in a flat-bottom tissue culture flask. Jurkat T cells were harvested when 70%–90% confluent and were resuspended in complete RPMI-1640 at a concentration of 100 000 cells/mL after centrifugation at 1500 g for 5 min at 4 °C. MBs were added to cell solutions at a 2.5% (v/v) concentration for calcein loading experiments unless otherwise indicated. For delivery efficiency assessment with different MB concentrations (0%–10% v/v), 10 μg/mL (1 mM) of 10 kDa fluorescein isothiocyanate (FITC)-dextran was utilized. Additionally, 5% (v/v) concentration was used for intracellular delivery of 26 kDa sfGFP. MBs with a cationic surface charge were used for all acoustofluidic experiments unless otherwise indicated. In separate experimental studies, calcein, 10 kDa FITC-Dextran, and 26 kDa sfGFP were added to Jurkat T cell/MB solutions at a final concentration of 100 μg/mL, 10 μg/mL(1 mM), and 52 μg/mL(2 μM), respectively. Unless otherwise noted, the molecular compound was added to each sample 1 min prior to acoustofluidic treatment. After treatment, each sample was washed three times via centrifugation with 3 ml of PBS to remove extracellular calcein prior to flow cytometry analysis. After washing each sample, propidium iodide (PI) was added 1 h post-acoustofluidic treatment to a final concentration of 20 μg/mL to detect non-viable cells via flow cytometry.

Flow cytometry was performed using a MACSQuant Analyzer 10 (Miltenyi Biotec, Auburn, CA) with 10 000 events recorded per sample. Data were analyzed using FlowJo software (Ashland, OR). Forward and side-scatter plots were gated to the main cell population (based on untreated control samples) and the mean fluorescence intensity in the FITC channel was calculated to assess intracellular calcein delivery. Cell viability was assessed by gating PI fluorescence intensity in the untreated control samples and determining the percentage of cells in the treatment groups that exceeded the gated PI fluorescence threshold using the PE/Texas Red channel.

Fluid dynamics in the acoustofluidic devices were modeled using ANSYS (Canonsburg, PA). Acoustofluidic channel geometries were imported from SolidWorks and transferred to the meshing component. Mesh sensitivity analysis was performed for element sizes between 0.2 and 3 mm to determine the optimal mesh sizes for each simulation. An element size of 0.5 mm (18 169 elements) was used for the 1-mm diameter rectilinear channel, an element size of 0.4 mm (49 495 elements) was used for the 2-mm diameter rectilinear channel, and an element size of 0.4 mm (30 690 elements) was used for the 1-mm diameter concentric spiral channel. Each mesh was transferred to ANSYS Fluent and a discrete phase model was used to compute the wall shear stress throughout each acoustofluidic channel, assuming laminar flow. To model perfluorobutane MBs within the acoustofluidic channels, particles were generated with a mean diameter of 2 μm, modeled as a Rosin-Rammler distribution, with a particle density of 24.6 kg/m3 and a specific heat of 809 J/(kg*K) in the discrete phase model. Given the very small MB diameter (2 μm) and the relatively short duration within the device (∼2–3 s), buoyancy is not expected to be a significant factor and was therefore ignored in the computational model. The expected terminal rise velocity of MBs in the acoustofluidic channel, Ub, was calculated assuming low Reynolds number as previously described (Stokes, 1851; Parmar and Majumder, 2015),

Ub=Db2gρlρg18μl,
(2)

where Db represents MB diameter, g represents acceleration due to gravity, ρl represents density of the liquid, ρg represents density of the gas, and μl represents viscosity of the liquid. Assuming a MB diameter of 2 μm, a liquid density of 1000 kg/m3 for water, a gas density of 11.2 kg/m3 for perfluorobutane, and a liquid viscosity of 1 mPa·s for water, the terminal rise velocity would be 2.1 μm/s. At this rate, the MB would be expected to rise no more than 6.3 μm, or 0.63% of the 1-mm channel diameter, while passing through the acoustofluidic device within 3 s. The fluid material was modeled as water at 25 °C (density of 1000 kg/m3 and viscosity of 0.91 cP). The channel inlet velocity was set to 0.002 m/s (1.5 ml/min) and the channel inlet pressure and outlet pressure boundary conditions were set to 1 atm for each simulation. The volume average wall shear stress was computed for each acoustofluidic channel geometry.

sfGFP (plasmid #51562, Addgene, Watertown, MA) was expressed in frame with a chitin-binding protein (CBP) and self-splicing intein protein spacer for purification purposes. The plasmid was transformed into the Escherichia coli strain BL21 Star (Thermo Fisher Scientific, Waltham, MA) and cells were grown on Luria-Bertani (LB) medium-based agarose plates containing 100 μg/mL ampicillin. Antibiotic resistant colonies were grown to an optical density of 0.6 using λ = 595 nm in liquid culture on an orbital shaker at 225 rpm and 37 °C in LB containing 100 μg/mL ampicillin. Protein expression was induced for 2 h after adding isopropyl-β-1-thiogalactopyranoside (IPTG) to a final concentration of 0.6 mM and the bacteria were harvested via centrifugation at 5000 g for 30 min at 4 °C. Cell pellets were resuspended in a small volume of buffer A (1 M NaCl, 50 mM Tris-HCl, pH 8.5) containing 1 mM phenylmethylsulfonyl fluoride (PMSF) to inhibit serine protease activity. The resuspended pellets were stored at −80 °C for up to four months before purification. For sfGFP purification, cells were lysed by sonication (Q500, Qsonica, Newtown, CT) and bacterial debris was removed by centrifugation for 30 min at 5000 g at 4 °C. The supernatant was loaded via gravity flow onto a 15 ml chitin resin (NEB Biolabs, Ipswich, MA) containing column. The column was then washed with 20 column volumes of buffer A. The sfGFP protein was eluted after incubation with 50 mM DTT dissolved in buffer A at 4 °C for 48 h. The eluted protein was then dialyzed into 50 mM phosphate buffer at a pH of 7.0 and concentrated using centrifugal filter units (Amicon Ultra 10 kDa, Millipore Sigma, St. Louis, MO). The purity of the protein was confirmed by SDS-PAGE and averaged at least 95%. Purified protein aliquots were snap frozen in liquid nitrogen and stored at −80 °C until used in experiments.

Statistical analysis was conducted using SPSS 26 (Armonk, NY). Statistical comparisons between groups were determined using analysis of variance (ANOVA), unless stated otherwise, with statistical significance defined as p < 0.05 (two-tailed). Post hoc analysis was conducted using Tukey's test for ANOVA. Bars represent mean ± standard deviations.

Ultrasound-mediated intracellular delivery of calcein to Jurkat T cells was evaluated with various MB phospholipid formulations at 2.5% (v/v) in PBS as measured by flow cytometry post-treatment. Control groups without ultrasound treatment were tested for each MB formulation to determine if MBs alone enhanced calcein delivery when samples passed through the fluidic channels without ultrasound exposure. In the absence of ultrasound exposure, there was no statistically significant difference in calcein delivery to T cells between samples containing MBs with neutral surface charge and samples containing MBs with a cationic surface charge. However, significant uptake of calcein was observed in Jurkat T cell samples containing MBs that were exposed to 3.8 MPa peak negative ultrasound pressure as they passed through the 1 mm rectilinear channel acoustofluidic device, compared to control groups without ultrasound treatment [Fig. 3(a), ANOVA p < 0.001, n = 5–6/group). Post hoc analysis indicated that calcein delivery was significantly enhanced in Jurkat T cell samples with cationic MBs that were treated with ultrasound compared to ultrasound-treated samples with neutral MBs or ultrasound-treated samples with no MBs (ANOVA p < 0.001).

FIG. 3.

(Color online) Effect of microbubble surface charge on molecular delivery in the 1 mm rectilinear channel acoustofluidic device. (A) Microbubbles with cationic surface charge enhanced calcein delivery to Jurkat T cells in the acoustofluidic device compared to microbubbles with neutral surface charge (ANOVA p < 0.001, n = 5–6/group). (B) Cell viability was moderately reduced after acoustofluidic treatment with cationic microbubbles compared to neutral microbubbles (ANOVA p < 0.001, n = 3/group). (C) There was no statistically significant difference in acoustic attenuation between cationic microbubbles and neutral microbubbles (p > 0.05, n = 15/group).

FIG. 3.

(Color online) Effect of microbubble surface charge on molecular delivery in the 1 mm rectilinear channel acoustofluidic device. (A) Microbubbles with cationic surface charge enhanced calcein delivery to Jurkat T cells in the acoustofluidic device compared to microbubbles with neutral surface charge (ANOVA p < 0.001, n = 5–6/group). (B) Cell viability was moderately reduced after acoustofluidic treatment with cationic microbubbles compared to neutral microbubbles (ANOVA p < 0.001, n = 3/group). (C) There was no statistically significant difference in acoustic attenuation between cationic microbubbles and neutral microbubbles (p > 0.05, n = 15/group).

Close modal

PI was added to samples post-treatment to assess cell viability as measured by flow cytometry. No statistically significant differences in cell viability were observed in ultrasound-treated samples with neutral MBs or no exogenous MBs compared to control groups without ultrasound treatment for each condition, respectively. However, post hoc analysis indicated that cell viability was significantly reduced in ultrasound-treated samples with cationic MBs after acoustofluidic treatment, compared to samples with cationic MBs which were not exposed to ultrasound as they passed through the acoustofluidic device [Fig. 3(b), ANOVA p = 0.001). These results indicate that MB surface charge has a significant impact on acoustofluidic-mediated effects in Jurkat T cells and cationic surface charge significantly enhances molecular delivery compared to neutral surface charge, but moderately reduced cell viability compared to respective control groups.

The impact of molecular exposure time points on intracellular delivery was investigated using the 1 mm rectilinear acoustofluidic device with ultrasound treatment at 3.8 MPa peak negative pressure output in Jurkat T cell samples with cationic MBs. The highest level of molecular delivery occurred when calcein was added 1 min prior to acoustofluidic treatment, with intracellular fluorescence enhanced by 6.26 ± 3.66-fold compared to baseline. Post hoc analysis with Tukey's test indicated that molecular delivery was significantly enhanced when samples were exposed to calcein at 1 min prior to acoustofluidic treatment compared to adding calcein at 1 min, 5 min, 15 min, or 30 min post-acoustofluidic treatment (Fig. 4, ANOVA p < 0.001, n = 5–6). These results indicate that biomolecule presence in solution during acoustofluidic treatment has a greater effect on molecular loading compared to biomolecule addition to solution post-treatment, which suggests that molecular delivery occurs at a faster rate during acoustofluidic treatment compared to other time points.

FIG. 4.

Effect of molecular exposure time point on calcein uptake in Jurkat T cells. Calcein was added to cell solutions at various timepoints before and after passing through the 1 mm rectilinear acoustofluidic device with ultrasound treatment (3.8 MPa peak negative pressure) and cationic microbubbles. Calcein delivery was enhanced when delivered prior to treatment compared to all post-treatment exposure time points (ANOVA p < 0.001, n = 5–6/group).

FIG. 4.

Effect of molecular exposure time point on calcein uptake in Jurkat T cells. Calcein was added to cell solutions at various timepoints before and after passing through the 1 mm rectilinear acoustofluidic device with ultrasound treatment (3.8 MPa peak negative pressure) and cationic microbubbles. Calcein delivery was enhanced when delivered prior to treatment compared to all post-treatment exposure time points (ANOVA p < 0.001, n = 5–6/group).

Close modal

The impact of acoustofluidic channel geometries on molecular delivery was investigated with three geometries (as shown in Fig. 1): a 1-mm diameter rectilinear channel, a 2-mm diameter rectilinear channel, and a concentric spiral design with a 1 mm channel diameter, using flow cytometry for analysis after acoustofluidic treatment. In all three channel geometries, ultrasound treatment significantly enhanced calcein delivery to Jurkat T cells compared to control groups without ultrasound [Fig. 5(a), ANOVA p < 0.001, n = 6/group]. Furthermore, post hoc analysis indicated that calcein delivery to Jurkat T cells was significantly enhanced in samples passed through the 1 mm concentric spiral channel with ultrasound treatment compared to ultrasound-treated samples passed through the 1 mm rectilinear channel (ANOVA p < 0.001) and 2 mm rectilinear channel (ANOVA p = 0.001), respectively.

FIG. 5.

(Color online) Effect of acoustofluidic channel geometry on molecular delivery. (A) Ultrasound treatment (3.8 MPa peak negative pressure) enhanced calcein delivery to Jurkat T cells in each channel geometry compared to control groups passed through the channels without ultrasound treatment (ANOVA p < 0.001, n = 6/group). With ultrasound treatment the 1 mm concentric spiral geometry significantly enhanced calcein delivery to Jurkat T cells compared to the 1 and 2 mm rectilinear channel geometry (ANOVA p < 0.001) and 2 mm rectilinear channel geometry (ANOVA p < 0.001). (B) Ultrasound treatment reduced cell viability in acoustofluidic devices (ANOVA p < 0.001, n = 3–6/group), but viability remained above 75% in all groups. (C) Ultrasound treatment reduced acoustic attenuation of microbubble solutions after treatment in each acoustofluidic channel geometry compared to the no flow control condition (ANOVA p < 0.001, n = 6/group), indicating that ultrasound treatment induced microbubble destruction in the channels. There was no statistically significant difference in acoustic attenuation between each channel geometry.

FIG. 5.

(Color online) Effect of acoustofluidic channel geometry on molecular delivery. (A) Ultrasound treatment (3.8 MPa peak negative pressure) enhanced calcein delivery to Jurkat T cells in each channel geometry compared to control groups passed through the channels without ultrasound treatment (ANOVA p < 0.001, n = 6/group). With ultrasound treatment the 1 mm concentric spiral geometry significantly enhanced calcein delivery to Jurkat T cells compared to the 1 and 2 mm rectilinear channel geometry (ANOVA p < 0.001) and 2 mm rectilinear channel geometry (ANOVA p < 0.001). (B) Ultrasound treatment reduced cell viability in acoustofluidic devices (ANOVA p < 0.001, n = 3–6/group), but viability remained above 75% in all groups. (C) Ultrasound treatment reduced acoustic attenuation of microbubble solutions after treatment in each acoustofluidic channel geometry compared to the no flow control condition (ANOVA p < 0.001, n = 6/group), indicating that ultrasound treatment induced microbubble destruction in the channels. There was no statistically significant difference in acoustic attenuation between each channel geometry.

Close modal

PI was added to samples post-treatment to assess cell viability as measured by flow cytometry. Reduced cell viability was observed with ultrasound treatment in acoustofluidic devices compared to control groups without ultrasound [Fig. 5(b), ANOVA p < 0.001, n = 3–6/group). Although reduced viability was observed, viability remained above 75% in all treatment groups. There was no statistically significant difference in viability between each channel geometry.

The effect of acoustofluidic channel geometry on ultrasound-induced MB destruction was also assessed using acoustic attenuation measurements. Ultrasound treatment significantly reduced acoustic attenuation in all geometries compared to the control group without ultrasound (ANOVA p < 0.001, n = 6/group). Post hoc analysis indicated that there was no statistically significant difference between acoustic attenuation after treatment in the 1 mm concentric spiral channel, 1 mm rectilinear channel, and 2 mm rectilinear channel, respectively.

Fluid dynamics within each acoustofluidic channel geometry were analyzed using computational modeling in ANSYS to further assess factors that may influence molecular delivery in each acoustofluidic channel geometry. The simulated Reynolds number in each channel geometry was below 28, indicating laminar flow within the acoustofluidic devices. As shown in Fig. 6, the average wall shear stress was tenfold higher in the 1-mm diameter rectilinear channel compared to the 2-mm diameter rectilinear channel, as would be expected given the relatively higher fluid velocity within the smaller channels. However, the average wall shear stress was even higher in the 1 mm concentric spiral channels (2.8-fold higher compared to the 1 mm rectilinear channel), indicating that the curvature of the concentric spiral channel geometry increases wall shear stress separately from fluid velocity and this may be a key factor that influences acoustofluidic molecular delivery to cells within the concentric spiral channel geometry.

FIG. 6.

(Color online) Computational modeling of wall shear stress in each acoustofluidic channel geometry for a flow rate of 1.5 ml/min. (A) Representative images of wall shear stress patterns from computational modeling of each acoustofluidic channel geometry. (B) Average wall shear stress was higher in the concentric spiral geometry compared to rectilinear geometries.

FIG. 6.

(Color online) Computational modeling of wall shear stress in each acoustofluidic channel geometry for a flow rate of 1.5 ml/min. (A) Representative images of wall shear stress patterns from computational modeling of each acoustofluidic channel geometry. (B) Average wall shear stress was higher in the concentric spiral geometry compared to rectilinear geometries.

Close modal

Acoustofluidic-mediated molecular delivery was evaluated at various ultrasound pressures using the 1-mm diameter concentric spiral design, as measured with flow cytometry. Acoustofluidic treatment was tested at a range of ultrasound pressures between 0 and 5.1 MPa peak negative ultrasound output pressure. There were no statistically significant differences in calcein delivery to Jurkat T cells at lower ultrasound output pressures (0–1.3 MPa, p > 0.05), whereas significant delivery was observed at ultrasound output pressures of 2.5 MPa and higher. Post hoc analysis indicated that calcein delivery was significantly enhanced at ultrasound output pressures of 2.5 MPa and higher compared to lower ultrasound output pressures [0–1.3 MPa; Fig. 7(a), ANOVA p < 0.001, n = 3/group]. Post hoc analysis also indicated that the highest acoustofluidic ultrasound pressure evaluated (5.1 MPa peak negative ultrasound output pressure) resulted in the highest amount of molecular delivery to Jurkat T cells compared to all other acoustofluidic ultrasound pressures tested in this study (ANOVA p < 0.001), with detected intracellular fluorescence intensity increased by a maximum of 22.09 ± 0.13-fold compared to the 0 MPa control group.

FIG. 7.

Effect of ultrasound pressure on molecular delivery to Jurkat T cells in acoustofluidic device. (A) The highest ultrasound pressure evaluated (5.1 MPa peak negative ultrasound output pressure) resulted in the highest amount of intracellular calcein delivery compared to all other treatment groups (ANOVA p < 0.001, n = 6/group). (B) Cell viability remained above 75% at all ultrasound pressure conditions (n = 3/group). (C) Acoustic attenuation was significantly reduced in samples treated at peak negative ultrasound pressures 3.8 MPa and higher, indicating ultrasound-mediated microbubble destruction at higher acoustic pressures (ANOVA p < 0.05, n = 5–7/group).

FIG. 7.

Effect of ultrasound pressure on molecular delivery to Jurkat T cells in acoustofluidic device. (A) The highest ultrasound pressure evaluated (5.1 MPa peak negative ultrasound output pressure) resulted in the highest amount of intracellular calcein delivery compared to all other treatment groups (ANOVA p < 0.001, n = 6/group). (B) Cell viability remained above 75% at all ultrasound pressure conditions (n = 3/group). (C) Acoustic attenuation was significantly reduced in samples treated at peak negative ultrasound pressures 3.8 MPa and higher, indicating ultrasound-mediated microbubble destruction at higher acoustic pressures (ANOVA p < 0.05, n = 5–7/group).

Close modal

Flow cytometry analysis of PI staining indicated a general trend that as acoustofluidic ultrasound pressures increased there was a corresponding decrease in cell viability, although this difference was relatively small and cell viability remained above 75% at all treatment conditions. These results indicate that acoustofluidic treatment can significantly enhance molecular delivery without significantly impacting cell viability.

Acoustic attenuation of each sample was measured to assess the effect of acoustofluidic ultrasound pressures on MB destruction. No evidence of MB destruction was observed at peak negative ultrasound output pressures of 0–2.5 MPa, but there was a significant decrease in acoustic attenuation after acoustofluidic treatment with ultrasound output pressures of 3.8 MPa and higher [Fig. 7(c); ANOVA p < 0.05, n = 5–7/group], indicating higher levels of MB destruction at these conditions. The general trend of acoustic attenuation as a function of ultrasound pressure was roughly inversely related to the trend in molecular delivery, with acoustic attenuation generally lower at ultrasound pressures where fluorescence intensity was higher. This suggests that MB destruction is associated with enhanced intracellular molecular delivery during acoustofluidic treatment.

Acoustofluidic treatment was tested with a range of cationic MB concentrations (0%–10% v/v) at 5.1 MPa peak negative ultrasound output pressure to assess molecular delivery with a large molecule (≥1 kDa). Molecular delivery was assessed using 10 kDa FITC-Dextran and a poor linear correlation with MB concentration was observed (R2 = 0.19) (see Fig. 8). This indicates that a MB concentration within a specific therapeutic window is required for optimal molecular delivery. A MB concentration of 5% (v/v) improved molecular delivery compared to lower MB concentrations (0%–1% v/v, ANOVA p < 0.001, n = 6–9/group) and higher MB concentrations (10% v/v, p < 0.01, n = 6–9).

FIG. 8.

(Color online) Intracellular delivery of 10 kDa FITC-dextran at various cationic microbubble concentrations, using a concentric spiral acoustofluidic device at 5.1 MPa peak negative ultrasound output pressure. (A) Acoustofluidic treatment with 10 μg/mL (1 mM) of 10 kDa FITC-Dextran increased intracellular delivery to Jurkat T cells at a microbubble concentration of 5% (v/v) compared to lower microbubble concentrations (0%–1% v/v, ANOVA p < 0.001, n = 6–9/group) and higher microbubble concentrations (10% v/v, p < 0.01). (B) Acoustofluidic treatment with a cationic microbubble concentration of 5% (v/v) had reduced cell viability compared to the 0% (v/v) microbubble concentration control group (ANOVA p < 0.001, n = 6–9/group), 1%–2.5% (v/v) microbubble concentrations (p < 0.05), and 10% (v/v) microbubble concentration (p < 0.05).

FIG. 8.

(Color online) Intracellular delivery of 10 kDa FITC-dextran at various cationic microbubble concentrations, using a concentric spiral acoustofluidic device at 5.1 MPa peak negative ultrasound output pressure. (A) Acoustofluidic treatment with 10 μg/mL (1 mM) of 10 kDa FITC-Dextran increased intracellular delivery to Jurkat T cells at a microbubble concentration of 5% (v/v) compared to lower microbubble concentrations (0%–1% v/v, ANOVA p < 0.001, n = 6–9/group) and higher microbubble concentrations (10% v/v, p < 0.01). (B) Acoustofluidic treatment with a cationic microbubble concentration of 5% (v/v) had reduced cell viability compared to the 0% (v/v) microbubble concentration control group (ANOVA p < 0.001, n = 6–9/group), 1%–2.5% (v/v) microbubble concentrations (p < 0.05), and 10% (v/v) microbubble concentration (p < 0.05).

Close modal
FIG. 9.

(Color online) Intracellular delivery of green fluorescent protein (GFP) using the concentric spiral acoustofluidic device at 5.1 MPa peak negative ultrasound output pressure and a cationic microbubble concentration of 5% (v/v). (A) Molecular delivery with 2 μM of 26 kDa sfGFP was increased after acoustofluidic treatment compared to flow only without ultrasound (negative control; Student's T-test p < 0.05, n = 6/group). (B) Cell viability was reduced after acoustofluidic treatment compared to flow only without ultrasound (Student's T-test p < 0.001, n = 6/group), but viability remained above 90% in each group.

FIG. 9.

(Color online) Intracellular delivery of green fluorescent protein (GFP) using the concentric spiral acoustofluidic device at 5.1 MPa peak negative ultrasound output pressure and a cationic microbubble concentration of 5% (v/v). (A) Molecular delivery with 2 μM of 26 kDa sfGFP was increased after acoustofluidic treatment compared to flow only without ultrasound (negative control; Student's T-test p < 0.05, n = 6/group). (B) Cell viability was reduced after acoustofluidic treatment compared to flow only without ultrasound (Student's T-test p < 0.001, n = 6/group), but viability remained above 90% in each group.

Close modal

Trypan blue assay indicated a weak negative linear correlation with MB concentration and cell viability (R2 = 0.17), which indicated that an increase in MB concentration is a poor predicator variable for cell viability. A MB concentration of 5% (v/v) reduced viability compared to flow only (0% v/v, ANOVA p < 0.01, n = 3/group) and 1% (v/v), 2.5% (v/v), 10% (v/v, ANOVA p < 0.05). Of significant interest, no correlation was found between molecular loading and cell viability (R2 = 0.50), suggesting that increased molecular delivery can be achieved without compromising cell viability.

A proteinaceous molecule, sfGFP (26 kDa), was utilized to assess acoustofluidic delivery efficiency at an ultrasound output pressure of 5.1 MPa and a MB concentration of 5% (v/v) in the concentric spiral acoustofluidic device (see Fig. 9). At an extracellular sgGFP concentration of 52 μg/mL (2 μM), a statistically significant increase in intracellular fluorescence was observed compared to the flow only control group without ultrasound treatment (1.86 ± 0.81-fold, Student's t-test p < 0.05, n = 6/group). Cell viability was reduced after acoustofluidic treatment compared to the flow only control group without ultrasound treatment (Student's t-test p < 0.001, n = 6/group), but viability remained above 90% in both groups.

Acoustofluidics is a rapidly developing field with a wide range of applications (Bruus et al., 2011; Gedge and Hill, 2012; Lenshof et al., 2012; Carugo et al., 2017; Wu et al., 2017; Belling et al., 2020). In this study, 3D-printed acoustofluidic devices were designed and evaluated to enhance molecular delivery to human T cells via MB-mediated mechanisms. Although single-element transducers with high duty cycle are often used for acoustofluidic applications, in this study we utilized a clinical ultrasound probe to transmit B-mode pulses into the acoustofluidic channels similar to our previous studies (Centner et al., 2020). The use of gas-filled MBs as ultrasound contrast agents reduces the amount of ultrasound energy necessary for molecular delivery compared to approaches without contrast agents. By adding MBs to cell solutions that pass through the 3D-printed acoustofluidic device, ultrasound-mediated biomolecular delivery to cells can be achieved rapidly as a single device can process volumes of at least 90 ml/h.

The results of this study demonstrate for the first time that the MB phospholipid surface charge and the molecular exposure time point both have important implications on biomolecular delivery to human T cells in acoustofluidic devices. The experimental findings in this study demonstrate several key parameters in the 3D-printed acoustofluidic devices that have significant effects on intracellular biomolecular delivery. Although this study focused on delivery of a small fluorescent molecule (calcein) for detection via flow cytometry, it is likely that optimization of the important parameters identified in this study will also enable enhanced delivery of many other biomolecules to cells using acoustofluidic devices, including proteins and nucleic acid molecules. Therefore, these findings provide new insights into the development of potential acoustofluidic techniques for manufacturing of cell-based therapies.

In this study, there was a statistically significant difference in molecular delivery based on the surface charge of the MB phospholipid coating [Fig. 2(a)], with the highest level of calcein delivery detected when cationic MBs were used with ultrasound treatment (3.8 MPa peak negative pressure output, p < 0.001 compared to control group without ultrasound treatment). It was previously determined that MB phospholipid surface charge could enhance ultrasound-mediated molecular delivery to HEK 293 cells (Tlaxca et al., 2010). However, there were distinct differences in the previous study by Tlaxca et al., including utilization of biotinylated-PEG substrate to conjugate the molecules, assessment using anchorage-dependent cells (unlike anchorage-independent Jurkat T cells), and the use of a static ultrasound configuration. These differences may account for the lack of differences in molecular delivery between cationic and neutral lipid-coated MBs when treated with ultrasound in other cell lines using the same ultrasound conditions (Tlaxca et al., 2010). It is likely that the positive surface charge on cationic MBs increases interactions with negatively-charged components of the T cell surface, such as the glycocalyx, via non-covalent charge-charge interactions (Springer, 1990). These interactions would reduce the distance between the MBs and cells, which may play a role in enhancing molecular delivery during ultrasound-driven MB activity, possibly through sonoporation mechanisms. However, variance in glycocalyx expression levels can influence MB-cell interaction. It is possible that low glycocalyx levels could significantly reduce molecular loading with cationic MBs. MBs with a neutral surface charge are also responsive to ultrasound waves as indicated by the acoustic attenuation measurements, however the amount of molecular delivery was significantly reduced when neutral MBs were used at equivalent ultrasound conditions compared to cationic MBs. This suggests that the neutral MBs were cavitating at a further distance from the cells and therefore had minimal effect on intracellular molecular delivery.

The effects of molecular exposure time point before and after acoustofluidic treatment was a critical factor in calcein delivery, with maximum delivery detected when calcein was added prior to acoustofluidic treatment. Intracellular calcein delivery was also enhanced, albeit to a lesser degree, when calcein was added at various timepoints after acoustofluidic treatment. One potential mechanism to explain this finding is an extended duration of membrane pore opening prior to resealing, but this appears to be unlikely given that pores typically reseal within less than 15 min after ultrasound treatment (Zhou et al., 2008; Helfield et al., 2016). A more likely explanation is the possibility that endocytosis activity may have increased after acoustofluidic treatment due to mechanical stress on cell membranes induced by ultrasound-driven MB activity (Juffermans et al., 2009). Although enhanced endocytosis activity may play some role, this does not appear to be the primary mode of intracellular delivery within the 3D-printed acoustofluidic channels given that much higher levels of molecular delivery were detected when calcein was added prior to acoustofluidic treatment.

This study also revealed that channel geometry is another key factor that influences acoustofluidic-enhanced molecular delivery, but this parameter had limited influence on cell viability. Acoustofluidic treatment with a concentric spiral channel geometry caused significantly higher levels of intracellular calcein delivery compared to acoustofluidic treatment with rectilinear channel geometries. The residence time of cells within the ultrasound beam is significantly increased in the concentric spiral channel geometry compared to rectilinear channel geometries, with a total channel length within the ultrasound beam of 57.7 mm for the concentric spiral geometry compared to 10 mm for each rectilinear geometry tested in this study. Residence time does not appear to have important implications on cell viability with the experimental parameters used in this study when comparing three different acoustofluidic device designs. Although residence time is likely a factor in acoustofluidic-enhanced molecular delivery, it does not appear to solely explain the differences in results between each channel geometry. Acoustofluidic treatment in rectilinear channels enhanced intracellular fluorescence by approximately threefold compared to control groups without ultrasound treatment. However, there was no statistically significant difference in acoustofluidic-mediated molecular delivery between the two rectilinear channel geometries that were tested in this study, despite the fact that the average fluid velocity would have been four times faster in the 1-mm channel compared to the 2-mm channel. If residence time was the only factor one would expect that the 2-mm rectilinear channel would have higher molecular delivery compared to the 1-mm rectilinear channel, due to the fourfold increase in residence time for cells within the ultrasound beam, but this was not observed. Therefore, other differences in fluid dynamics between channel geometries may also be a key factor. Indeed, computational analysis of wall shear stress for each channel geometry revealed that the concentric spiral had significantly higher average wall shear stress compared to the rectilinear channels, even when the channel diameters were identical. The centrifugal force in spiral geometries creates accelerated velocity near the outer wall, resulting in a sharper velocity gradient and, hence, higher wall shear stress. These results suggest that the curvature of the acoustofluidic channel may also be a key factor in molecular delivery efficiency.

Additionally, the effect of ultrasound pressure amplitude was also found to play a key role on enhancing intracellular molecular delivery. As shown in Fig. 5, higher ultrasound pressures (2.4–5.1 MPa peak negative ultrasound pressure output) showed enhanced delivery of calcein compared to lower ultrasound pressures (0–1.4 MPa peak negative ultrasound pressure output). This indicates that molecular delivery is significantly enhanced above a specific ultrasound pressure threshold. Fundamental acoustic parameters that influence molecular delivery in anchorage independent cells have yet to be fully characterized, as multiple modalities could influence molecular delivery, including thermal and mechanical bioeffects. Thermal bioeffects have been observed with ultrasound-based platforms (Haar and Coussios, 2007; Bruningk et al., 2019); however, thermal effects were reduced in this acoustofluidic system by utilization of short B-mode pulses (2 μs) with a low duty cycle (below 1%) and short residence time (∼2–3 s in concentric spiral), which is more likely to be conducive to mechanical bioeffects over thermal bioeffects (Haar and Coussios, 2007). Mechanical bioeffects, specifically acoustic radiation force, could influence cell-MB interaction by displacing cells and particles to certain regions of the acoustofluidic device (Bruus et al., 2011; Bruus, 2012). Previous studies have shown the capacity for low pressure, pulsed ultrasound to facilitate displacement of ultrasound contrast agents toward vessel walls (Dayton et al., 1999a; Dayton et al., 1999b; Zhao et al., 2004). This could help facilitate molecular delivery to cells via cell-MB interactions and migrating towards vessel walls where shear stress is higher. Shear stress has been shown to facilitate molecular delivery in acoustofluidic systems (Belling et al., 2020). However, the effect of shear stress on molecular delivery with higher acoustic pressures and MBs is not fully understood and requires further investigation. Furthermore, our study utilized short ultrasound pulses (duty cycle less than 1%), which may reduce acoustic radiation force effects within the acoustofluidic channel compared to high duty cycles. Prior studies in static chambers have demonstrated that stable or inertial cavitation activity can enhance molecular delivery via sonoporation or other mechanisms (Paula et al., 2011; Fan et al., 2014; Delalande et al., 2015; Helfield et al., 2016). Although MB cavitation activity was not directly measured in this study, there was a significant decrease in acoustic attenuation after acoustofluidic treatment, which indicates MBs were destroyed during ultrasound exposure at higher pressures, as would be expected with higher levels of cavitation activity. Additionally, increased ultrasound pressures have been shown to play a critical role in molecular delivery via stable cavitation and inertial cavitation effects (Wu and Nyborg, 2008). For molecular delivery applications to T cells, inertial cavitation may be required for sufficient delivery, which requires higher acoustic pressures. Although the specific mechanisms of action have not been fully elucidated, this study provides evidence that ultrasound-driven MB activity has a key role in acoustofluidic-enhanced molecular delivery to cells.

Additionally, MB concentration influenced molecular delivery and cell viability. A weak relationship was observed between MB concentration and molecular delivery (R2 = 0.19), even with a significant (5.1 MPa) peak negative ultrasound output pressure. Interestingly, the highest MB concentration (10% v/v) was associated with reduced molecular delivery while maintaining higher cell viability. It is typically presumed that an increased MB concentration corresponds to increased molecular delivery while potentially inducing more cytotoxic effects. In this circumstance, it is feasible that a high MB concentration (≥10% v/v) could induce high levels of ultrasound multiple scattering within the acoustofluidic channels and reduce the acoustic pressure in situ. A similar trend has also been reported in a previous acoustofluidic study, albeit with a different cell type (erythrocytes), a different device design (polydimethylsiloxane), and different acoustic parameters with a lower peak negative output pressure (Centner et al., 2020). In this study, the ultrasound frequency was aligned with the MB resonance frequency in an attempt to induce sufficient MB oscillation for MB destruction and intracellular molecular delivery (Doinikov et al., 2009). However, at high MB concentrations, there may be significant scattering of ultrasound waves which reduces acoustic pressures within most of the acoustofluidic channel and causes less oscillation. This may be particularly critical with short ultrasound pulses in an acoustofluidic device where the MBs and cells are typically exposed to ultrasound waves for only ∼2–3 s.

This study explored ultrasound-enhanced delivery of a small molecule (i.e., calcein) and large molecules (i.e., 10 kDa FITC-Dextran and 26 kDa sfGFP) to human T cells in acoustofluidic channels to assess feasibility for potential cell-based therapy applications, such as CAR-T. The selected molecules are highly fluorescent and do not readily cross-cell membranes, thus they are useful for assessing intracellular molecular delivery via flow cytometry. In our study, we demonstrated that the delivery efficiency of larger biomolecules (molecular weight ≥10 kDa) could be achieved with an acoustofluidic device, which is necessary for proteinaceous or nucleic acid intracellular delivery. With 1 mM of 10 kDa FITC-Dextran in solution, over a 20-fold increase in intracellular fluorescence was observed at MB concentrations between 2.5% and 5% (v/v), compared to the flow only control group without ultrasound treatment. Intracellular fluorescence after acoustofluidic treatment with sfGFP was only 1.86 ± 0.81-fold higher than the flow only control group, but this lower amount of sfGFP delivery compared to FITC-dextran delivery is not surprising given the 500× difference in extracellular molarity between sfGFP (2 μM) and 10 kDa FITC-Dextran (1 mM) in this study. The extracellular fluorophore concentration could play a significant role in delivery efficiency as we have previously described (Bhutto et al., 2018). Even though significant molecular loading was achieved with 10 kDa FITC-Dextran, it is likely lower compared to small molecules such as calcein (molecular weight of 0.6 kDa), but may still be sufficient to provide the functional effects that are needed for therapeutic efficacy (Bhutto et al., 2018). Further studies are required to determine required cell transfection efficiency for therapeutic applications with an acoustofluidic device. A multitude of factors likely will influence efficiency of molecular delivery, including cell type, donor source, required cell dose, and cell viability. In addition, multiple doses may also increase molecular delivery to T cells. These factors will likely allow optimization of the acoustofluidic platform to enable manipulation of cells via CRISPR-cas9 gene editing or other techniques that are often used for cell therapy manufacturing applications.

The results of this study demonstrate the feasibility of utilizing a 3D-printed acoustofluidic device to rapidly load biomolecular compounds into human Jurkat T cells. Additionally, this study identifies several key acoustic and fluidic parameters that significantly influence the efficiency of intracellular delivery, including MB surface charge, MB concentration, the time point when the molecular compound is added to solution relative to the acoustofluidic treatment time point, flow channel geometries, and ultrasound pressure. These results provide new insights into techniques to optimize acoustofluidic delivery of biomolecules to human T cells. Further development and validation of acoustofluidic platforms may enable utilization of this technique to improve manufacturing of cell-based therapies, such as CAR-T immunotherapies for treatment of cancer and other diseases.

This research was funded by an NSF Partnership for Innovation Grant (Award No. 1827521). The Flow Cytometry Core at the University of Louisville Christina Lee Brown Envirome Institute provided technical assistance. We thank Sankar Adhya and Francisco Malagon for the gracious donation of TXB1-sfGFP vector (Addgene plasmid No. 51562: http://n2t.net/addgene:51562; RRID:Addgene_51562). Co-authors M.A.M. and J.A.K. are co-inventors on intellectual property related to this research.

1.
Annesley
,
C. E.
,
Summers
,
C.
,
Ceppi
,
F.
, and
Gardner
,
R. A.
(
2018
). “
The evolution and future of CAR T cells for B-cell acute lymphoblastic leukemia
,”
Clin. Pharmacol. Ther.
103
,
591
598
.
2.
Belling
,
J. N.
,
Heidenreich
,
L. K.
,
Tian
,
Z.
,
Mendoza
,
A. M.
,
Chiou
,
T. T.
,
Gong
,
Y.
,
Chen
,
N. Y.
,
Young
,
T. D.
,
Wattanatorn
,
N.
,
Park
,
J. H.
,
Scarabelli
,
L.
,
Chiang
,
N.
,
Takahashi
,
J.
,
Young
,
S. G.
,
Stieg
,
A. Z.
,
De Oliveira
,
S.
,
Huang
,
T. J.
,
Weiss
,
P. S.
, and
Jonas
,
S. J.
(
2020
). “
Acoustofluidic sonoporation for gene delivery to human hematopoietic stem and progenitor cells
,”
Proc. Natl. Acad. Sci. U.S.A.
117
,
10976
10982
.
3.
Bhutto
,
D. F.
,
Murphy
,
E. M.
,
Priddy
,
M. C.
,
Centner
,
C. C.
,
Moore Iv
,
J. B.
,
Bolli
,
R.
, and
Kopechek
,
J. A.
(
2018
). “
Effect of molecular weight on sonoporation-mediated uptake in human cells
,”
Ultrasound Med. Biol.
44
,
2662
2672
.
4.
Bruningk
,
S. C.
,
Rivens
,
I.
,
Mouratidis
,
P.
, and
Ter Haar
,
G.
(
2019
). “
Focused ultrasound-mediated hyperthermia in vitro: An experimental arrangement for treating cells under tissue-mimicking conditions
,”
Ultrasound Med. Biol.
45
,
3290
3297
.
5.
Bruus
,
H.
(
2012
). “
Acoustofluidics 7: The acoustic radiation force on small particles
,”
Lab Chip
12
,
1014
1021
.
6.
Bruus
,
H.
,
Dual
,
J.
,
Hawkes
,
J.
,
Hill
,
M.
,
Laurell
,
T.
,
Nilsson
,
J.
,
Radel
,
S.
,
Sadhal
,
S.
, and
Wiklund
,
M.
(
2011
). “
Forthcoming Lab on a Chip tutorial series on acoustofluidics: Acoustofluidics-exploiting ultrasonic standing wave forces and acoustic streaming in microfluidic systems for cell and particle manipulation
,”
Lab Chip
11
,
3579
3580
.
7.
Buzhor
,
E.
,
Leshansky
,
L.
,
Blumenthal
,
J.
,
Barash
,
H.
,
Warshawsky
,
D.
,
Mazor
,
Y.
, and
Shtrichman
,
R.
(
2014
). “
Cell-based therapy approaches: The hope for incurable diseases
,”
Regen. Med.
9
,
649
672
.
8.
Carugo
,
D.
,
Ankrett
,
D. N.
,
Glynne-Jones
,
P.
,
Capretto
,
L.
,
Boltryk
,
R. J.
,
Zhang
,
X.
,
Townsend
,
P. A.
, and
Hill
,
M.
(
2011
). “
Contrast agent-free sonoporation: The use of an ultrasonic standing wave microfluidic system for the delivery of pharmaceutical agents
,”
Biomicrofluidics
5
,
044108
4410815
.
9.
Carugo
,
D.
,
Aron
,
M.
,
Sezgin
,
E.
,
Bernardino de la Serna
,
J.
,
Kuimova
,
M. K.
,
Eggeling
,
C.
, and
Stride
,
E.
(
2017
). “
Modulation of the molecular arrangement in artificial and biological membranes by phospholipid-shelled microbubbles
,”
Biomaterials
113
,
105
117
.
10.
Centner
,
C. S.
,
Murphy
,
E. M.
,
Priddy
,
M. C.
,
Moore
,
J. T.
,
Janis
,
B. R.
,
Menze
,
M. A.
,
DeFilippis
,
A. P.
, and
Kopechek
,
J. A.
(
2020
). “
Ultrasound-induced molecular delivery to erythrocytes using a microfluidic system
,”
Biomicrofluidics
14
,
024114
.
11.
Chen
,
Y.
,
Li
,
S.
,
Gu
,
Y.
,
Li
,
P.
,
Ding
,
X.
,
Wang
,
L.
,
McCoy
,
J. P.
,
Levine
,
S. J.
, and
Huang
,
T. J.
(
2014
). “
Continuous enrichment of low-abundance cell samples using standing surface acoustic waves (SSAW)
,”
Lab Chip
14
,
924
930
.
12.
Dayton
,
P.
,
Klibanov
,
A.
,
Brandenburger
,
G.
, and
Ferrara
,
K.
(
1999a
). “
Acoustic radiation force in vivo: A mechanism to assist targeting of microbubbles
,”
Ultrasound Med. Biol.
25
,
1195
1201
.
13.
Dayton
,
P. A.
,
Morgan
,
K. E.
,
Klibanov
,
A. L.
,
Brandenburger
,
G. H.
, and
Ferrara
,
K. W.
(
1999b
). “
Optical and acoustical observations of the effects of ultrasound on contrast agents
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
46
,
220
232
.
14.
Deeks
,
S. G.
,
Wagner
,
B.
,
Anton
,
P. A.
,
Mitsuyasu
,
R. T.
,
Scadden
,
D. T.
,
Huang
,
C.
,
Macken
,
C.
,
Richman
,
D. D.
,
Christopherson
,
C.
,
June
,
C. H.
,
Lazar
,
R.
,
Broad
,
D. F.
,
Jalali
,
S.
, and
Hege
,
K. M.
(
2002
). “
A phase II randomized study of HIV-specific T-cell gene therapy in subjects with undetectable plasma viremia on combination antiretroviral therapy
,”
Mol. Ther.
5
,
788
797
.
15.
Delalande
,
A.
,
Leduc
,
C.
,
Midoux
,
P.
,
Postema
,
M.
, and
Pichon
,
C.
(
2015
). “
Efficient gene delivery by sonoporation is associated with microbubble entry into cells and the clathrin-dependent endocytosis pathway
,”
Ultrasound Med. Biol.
41
,
1913
1926
.
16.
Doinikov
,
A. A.
,
Haac
,
J. F.
, and
Dayton
,
P. A.
(
2009
). “
Resonance frequencies of lipid-shelled microbubbles in the regime of nonlinear oscillations
,”
Ultrasonics
49
,
263
268
.
17.
Falk
,
H.
,
Forde
,
P. F.
,
Bay
,
M. L.
,
Mangalanathan
,
U. M.
,
Hojman
,
P.
,
Soden
,
D. M.
, and
Gehl
,
J.
(
2017
). “
Calcium electroporation induces tumor eradication, long-lasting immunity and cytokine responses in the CT26 colon cancer mouse model
,”
Oncoimmunology
6
,
e1301332
.
18.
Fan
,
Z.
,
Kumon
,
R. E.
, and
Deng
,
C. X.
(
2014
). “
Mechanisms of microbubble-facilitated sonoporation for drug and gene delivery
,”
Ther. Deliv.
5
,
467
486
.
19.
Friend
,
J.
, and
Yeo
,
L. Y.
(
2011
). “
Microscale acoustofluidics: Microfluidics driven via acoustics and ultrasonics
,”
Rev. Mod. Phys.
83
,
647
704
.
20.
Gedge
,
M.
, and
Hill
,
M.
(
2012
). “
Acoustofluidics 17: Theory and applications of surface acoustic wave devices for particle manipulation
,”
Lab Chip
12
,
2998
3007
.
21.
Ghani
,
K.
,
Wang
,
X.
,
de Campos-Lima
,
P. O.
,
Olszewska
,
M.
,
Kamen
,
A.
,
Riviere
,
I.
, and
Caruso
,
M.
(
2009
). “
Efficient human hematopoietic cell transduction using RD114- and GALV-pseudotyped retroviral vectors produced in suspension and serum-free media
,”
Hum. Gene. Ther.
20
,
966
974
.
22.
Haar
,
G. T.
, and
Coussios
,
C.
(
2007
). “
High intensity focused ultrasound: Physical principles and devices
,”
Int. J. Hyperthermia
23
,
89
104
.
23.
He
,
C.
,
Wang
,
J.
,
Sun
,
S.
,
Zhang
,
Y.
, and
Li
,
S.
(
2019
). “
Immunomodulatory effect after irreversible electroporation in patients with locally advanced pancreatic cancer
,”
J. Oncol.
2019
,
9346017
.
24.
Helfield
,
B.
,
Chen
,
X.
,
Watkins
,
S. C.
, and
Villanueva
,
F. S.
(
2016
). “
Biophysical insight into mechanisms of sonoporation
,”
Proc. Natl. Acad. Sci. U.S.A.
113
,
9983
9988
.
25.
Hirsch
,
M. L.
,
Wolf
,
S. J.
, and
Samulski
,
R. J.
(
2016
). “
Delivering Transgenic DNA exceeding the carrying capacity of AAV vectors
,”
Methods Mol. Biol.
1382
,
21
39
.
26.
Juffermans
,
L. J.
,
Meijering
,
D. B.
,
van Wamel
,
A.
,
Henning
,
R. H.
,
Kooiman
,
K.
,
Emmer
,
M.
,
de Jong
,
N.
,
van Gilst
,
W. H.
,
Musters
,
R.
,
Paulus
,
W. J.
,
van Rossum
,
A. C.
,
Deelman
,
L. E.
, and
Kamp
,
O.
(
2009
). “
Ultrasound and microbubble-targeted delivery of therapeutic compounds: ICIN Report Project 49: Drug and gene delivery through ultrasound and microbubbles
,”
Neth. Heart J.
17
,
82
86
.
27.
Kim
,
K.
, and
Lee
,
W. G.
(
2017
). “
Electroporation for nanomedicine: A review
,”
J. Mater. Chem. B
5
,
2726
2738
.
28.
Kochenderfer
,
J. N.
,
Dudley
,
M. E.
,
Kassim
,
S. H.
,
Somerville
,
R. P.
,
Carpenter
,
R. O.
,
Stetler-Stevenson
,
M.
,
Yang
,
J. C.
,
Phan
,
G. Q.
,
Hughes
,
M. S.
,
Sherry
,
R. M.
,
Raffeld
,
M.
,
Feldman
,
S.
,
Lu
,
L.
,
Li
,
Y. F.
,
Ngo
,
L. T.
,
Goy
,
A.
,
Feldman
,
T.
,
Spaner
,
D. E.
,
Wang
,
M. L.
,
Chen
,
C. C.
,
Kranick
,
S. M.
,
Nath
,
A.
,
Nathan
,
D. A.
,
Morton
,
K. E.
,
Toomey
,
M. A.
, and
Rosenberg
,
S. A.
(
2015
). “
Chemotherapy-refractory diffuse large B-cell lymphoma and indolent B-cell malignancies can be effectively treated with autologous T cells expressing an anti-CD19 chimeric antigen receptor
,”
J. Clin. Oncol.
33
,
540
549
.
29.
Kochenderfer
,
J. N.
,
Somerville
,
R. P. T.
,
Lu
,
T.
,
Shi
,
V.
,
Bot
,
A.
,
Rossi
,
J.
,
Xue
,
A.
,
Goff
,
S. L.
,
Yang
,
J. C.
,
Sherry
,
R. M.
,
Klebanoff
,
C. A.
,
Kammula
,
U. S.
,
Sherman
,
M.
,
Perez
,
A.
,
Yuan
,
C. M.
,
Feldman
,
T.
,
Friedberg
,
J. W.
,
Roschewski
,
M. J.
,
Feldman
,
S. A.
,
McIntyre
,
L.
,
Toomey
,
M. A.
, and
Rosenberg
,
S. A.
(
2017
). “
Lymphoma remissions caused by anti-CD19 chimeric antigen receptor T cells are associated with high serum interleukin-15 levels
,”
J. Clin. Oncol.
35
,
1803
1813
.
30.
Kopechek
,
J. A.
,
Carson
,
A. R.
,
McTiernan
,
C. F.
,
Chen
,
X.
,
Hasjim
,
B.
,
Lavery
,
L.
,
Sen
,
M.
,
Grandis
,
J. R.
, and
Villanueva
,
F. S.
(
2015
). “
Ultrasound targeted microbubble destruction-mediated delivery of a transcription factor decoy inhibits STAT3 signaling and tumor growth
,”
Theranostics
5
,
1378
1387
.
31.
Kopechek
,
J. A.
,
Carson
,
A. R.
,
McTiernan
,
C. F.
,
Chen
,
X.
,
Klein
,
E. C.
, and
Villanueva
,
F. S.
(
2016
). “
Cardiac gene expression knockdown using small inhibitory RNA-loaded microbubbles and ultrasound
,”
PLoS One
11
,
e0159751
.
32.
Kopechek
,
J. A.
,
Haworth
,
K. J.
,
Raymond
,
J. L.
,
Douglas Mast
,
T.
,
Perrin
,
S. R.
,
Klegerman
,
M. E.
,
Huang
,
S.
,
Porter
,
T. M.
,
McPherson
,
D. D.
, and
Holland
,
C. K.
(
2011
). “
Acoustic characterization of echogenic liposomes: Frequency-dependent attenuation and backscatter
,”
J. Acoust. Soc. Am.
130
,
3472
3481
.
33.
Kopechek
,
J. A.
,
McTiernan
,
C. F.
,
Chen
,
X.
,
Zhu
,
J.
,
Mburu
,
M.
,
Feroze
,
R.
,
Whitehurst
,
D. A.
,
Lavery
,
L.
,
Cyriac
,
J.
, and
Villanueva
,
F. S.
(
2019
). “
Ultrasound and microbubble-targeted delivery of a microRNA Inhibitor to the heart suppresses cardiac hypertrophy and preserves cardiac function
,”
Theranostics
9
,
7088
7098
.
34.
Kurashina
,
Y.
,
Takemura
,
K.
, and
Friend
,
J.
(
2017
). “
Cell agglomeration in the wells of a 24-well plate using acoustic streaming
,”
Lab Chip
17
,
876
886
.
35.
Lenshof
,
A.
,
Magnusson
,
C.
, and
Laurell
,
T.
(
2012
). “
Acoustofluidics 8: Applications of acoustophoresis in continuous flow microsystems
,”
Lab Chip
12
,
1210
1223
.
36.
Li
,
H.
,
Friend
,
J. R.
, and
Yeo
,
L. Y.
(
2007
). “
Surface acoustic wave concentration of particle and bioparticle suspensions
,”
Biomed. Microdev.
9
,
647
656
.
37.
Majumder
,
J.
,
Taratula
,
O.
, and
Minko
,
T.
(
2019
). “
Nanocarrier-based systems for targeted and site specific therapeutic delivery
,”
Adv. Drug Deliv. Rev.
144
,
57
77
.
38.
Maude
,
S. L.
,
Laetsch
,
T. W.
,
Buechner
,
J.
,
Rives
,
S.
,
Boyer
,
M.
,
Bittencourt
,
H.
,
Bader
,
P.
,
Verneris
,
M. R.
,
Stefanski
,
H. E.
,
Myers
,
G. D.
,
Qayed
,
M.
,
De Moerloose
,
B.
,
Hiramatsu
,
H.
,
Schlis
,
K.
,
Davis
,
K. L.
,
Martin
,
P. L.
,
Nemecek
,
E. R.
,
Yanik
,
G. A.
,
Peters
,
C.
,
Baruchel
,
A.
,
Boissel
,
N.
,
Mechinaud
,
F.
,
Balduzzi
,
A.
,
Krueger
,
J.
,
June
,
C. H.
,
Levine
,
B. L.
,
Wood
,
P.
,
Taran
,
T.
,
Leung
,
M.
,
Mueller
,
K. T.
,
Zhang
,
Y.
,
Sen
,
K.
,
Lebwohl
,
D.
,
Pulsipher
,
M. A.
, and
Grupp
,
S. A.
(
2018
). “
Tisagenlecleucel in children and young adults with B-cell lymphoblastic leukemia
,”
N. Engl. J. Med.
378
,
439
448
.
39.
Miller
,
D. L.
,
Bao
,
S.
, and
Morris
,
J. E.
(
1999
). “
Sonoporation of cultured cells in the rotating tube exposure system
,”
Ultrasound Med. Biol.
25
,
143
149
.
40.
Moghimi
,
B.
,
Muthugounder
,
S.
,
Jambon
,
S.
,
Tibbetts
,
R.
,
Hung
,
L.
,
Bassiri
,
H.
,
Hogarty
,
M. D.
,
Barrett
,
D. M.
,
Shimada
,
H.
, and
Asgharzadeh
,
S.
(
2021
). “
Preclinical assessment of the efficacy and specificity of GD2-B7H3 SynNotch CAR-T in metastatic neuroblastoma
,”
Nat. Commun.
12
,
511
.
41.
Morgan
,
R. A.
,
Gray
,
D.
,
Lomova
,
A.
, and
Kohn
,
D. B.
(
2017
). “
Hematopoietic stem cell gene therapy: Progress and lessons learned
,”
Cell Stem Cell
21
,
574
590
.
42.
Myrset
,
A. H.
,
Fjerdingstad
,
H. B.
,
Bendiksen
,
R.
,
Arbo
,
B. E.
,
Bjerke
,
R. M.
,
Johansen
,
J. H.
,
Kulseth
,
M. A.
, and
Skurtveit
,
R.
(
2011
). “
Design and characterization of targeted ultrasound microbubbles for diagnostic use
,”
Ultrasound Med. Biol.
37
,
136
150
.
43.
Oh
,
N.
, and
Park
,
J. H.
(
2014
). “
Endocytosis and exocytosis of nanoparticles in mammalian cells
,”
Int. J. Nanomed.
9
,
51
63
.
44.
Parmar
,
R.
, and
Majumder
,
S. K.
(
2015
). “
Terminal rise velocity, size distribution and stability of microbubble suspension
,”
Asia-Pac. J. Chem. Eng.
10
,
450
465
.
45.
Paula
,
D. M.
,
Valero-Lapchik
,
V. B.
,
Paredes-Gamero
,
E. J.
, and
Han
,
S. W.
(
2011
). “
Therapeutic ultrasound promotes plasmid DNA uptake by clathrin-mediated endocytosis
,”
J. Gene Med.
13
,
392
401
.
46.
Pereno
,
V.
,
Aron
,
M.
,
Vince
,
O.
,
Mannaris
,
C.
,
Seth
,
A.
,
de Saint Victor
,
M.
,
Lajoinie
,
G.
,
Versluis
,
M.
,
Coussios
,
C.
,
Carugo
,
D.
, and
Stride
,
E.
(
2018
). “
Layered acoustofluidic resonators for the simultaneous optical and acoustic characterisation of cavitation dynamics, microstreaming, and biological effects
,”
Biomicrofluidics
12
,
034109
.
47.
Petersson
,
F.
,
Aberg
,
L.
,
Sward-Nilsson
,
A. M.
, and
Laurell
,
T.
(
2007
). “
Free flow acoustophoresis: Microfluidic-based mode of particle and cell separation
,”
Anal. Chem.
79
,
5117
5123
.
48.
Piscopo
,
N. J.
,
Mueller
,
K. P.
,
Das
,
A.
,
Hematti
,
P.
,
Murphy
,
W. L.
,
Palecek
,
S. P.
,
Capitini
,
C. M.
, and
Saha
,
K.
(
2018
). “
Bioengineering solutions for manufacturing challenges in CAR T cells
,”
Biotechnol. J.
13
,
1700095
.
49.
Sahay
,
G.
,
Querbes
,
W.
,
Alabi
,
C.
,
Eltoukhy
,
A.
,
Sarkar
,
S.
,
Zurenko
,
C.
,
Karagiannis
,
E.
,
Love
,
K.
,
Chen
,
D.
,
Zoncu
,
R.
,
Buganim
,
Y.
,
Schroeder
,
A.
,
Langer
,
R.
, and
Anderson
,
D. G.
(
2013
). “
Efficiency of siRNA delivery by lipid nanoparticles is limited by endocytic recycling
,”
Nat. Biotechnol.
31
,
653
658
.
50.
Schuster
,
S. J.
,
Svoboda
,
J.
,
Chong
,
E. A.
,
Nasta
,
S. D.
,
Mato
,
A. R.
,
Anak
,
O.
,
Brogdon
,
J. L.
,
Pruteanu-Malinici
,
I.
,
Bhoj
,
V.
,
Landsburg
,
D.
,
Wasik
,
M.
,
Levine
,
B. L.
,
Lacey
,
S. F.
,
Melenhorst
,
J. J.
,
Porter
,
D. L.
, and
June
,
C. H.
(
2017
). “
Chimeric antigen receptor T cells in refractory B-cell lymphomas
,”
N. Engl. J. Med.
377
,
2545
2554
.
51.
Sharei
,
A.
,
Zoldan
,
J.
,
Adamo
,
A.
,
Sim
,
W. Y.
,
Cho
,
N.
,
Jackson
,
E.
,
Mao
,
S.
,
Schneider
,
S.
,
Han
,
M. J.
,
Lytton-Jean
,
A.
,
Basto
,
P. A.
,
Jhunjhunwala
,
S.
,
Lee
,
J.
,
Heller
,
D. A.
,
Kang
,
J. W.
,
Hartoularos
,
G. C.
,
Kim
,
K. S.
,
Anderson
,
D. G.
,
Langer
,
R.
, and
Jensen
,
K. F.
(
2013
). “
A vector-free microfluidic platform for intracellular delivery
,”
Proc. Natl. Acad. Sci. U.S.A.
110
,
2082
2087
.
52.
Shi
,
J.
,
Huang
,
H.
,
Stratton
,
Z.
,
Huang
,
Y.
, and
Huang
,
T. J.
(
2009
). “
Continuous particle separation in a microfluidic channel via standing surface acoustic waves (SSAW)
,”
Lab Chip
9
,
3354
3359
.
53.
Shilton
,
R.
,
Tan
,
M. K.
,
Yeo
,
L. Y.
, and
Friend
,
J.
(
2008
). “
Particle concentration and mixing in microdrops driven by focused surface acoustic waves
,”
J. Appl. Phys.
104
,
014910
.
54.
Springer
,
T. A.
(
1990
). “
Adhesion receptors of the immune system
,”
Nature
346
,
425
434
.
55.
Stokes
,
G. G.
(
1851
). “
On the effect of the internal friction of fluids on the motion of pendulums
,”
Trans. Cambridge Philosoph. Soc.
9
,
8
.
56.
Tlaxca
,
J. L.
,
Anderson
,
C. R.
,
Klibanov
,
A. L.
,
Lowrey
,
B.
,
Hossack
,
J. A.
,
Alexander
,
J. S.
,
Lawrence
,
M. B.
, and
Rychak
,
J. J.
(
2010
). “
Analysis of in vitro transfection by sonoporation using cationic and neutral microbubbles
,”
Ultrasound Med. Biol.
36
,
1907
1918
.
57.
Turtle
,
C. J.
,
Hay
,
K. A.
,
Hanafi
,
L. A.
,
Li
,
D.
,
Cherian
,
S.
,
Chen
,
X.
,
Wood
,
B.
,
Lozanski
,
A.
,
Byrd
,
J. C.
,
Heimfeld
,
S.
,
Riddell
,
S. R.
, and
Maloney
,
D. G.
(
2017
). “
Durable molecular remissions in chronic lymphocytic leukemia treated with CD19-specific chimeric antigen receptor-modified T cells after failure of ibrutinib
,”
J. Clin. Oncol.
35
,
3010
3020
.
58.
Versluis
,
M.
,
Stride
,
E.
,
Lajoinie
,
G.
,
Dollet
,
B.
, and
Segers
,
T.
(
2020
). “
Ultrasound contrast agent modeling: A review
,”
Ultrasound Med. Biol.
46
,
2117
2144
.
59.
Wang
,
X.
, and
Riviere
,
I.
(
2015
). “
Manufacture of tumor- and virus-specific T lymphocytes for adoptive cell therapies
,”
Cancer Gene Ther.
22
,
85
94
.
60.
Wang
,
Z.
,
Li
,
F.
,
Rufo
,
J.
,
Chen
,
C.
,
Yang
,
S.
,
Li
,
L.
,
Zhang
,
J.
,
Cheng
,
J.
,
Kim
,
Y.
,
Wu
,
M.
,
Abemayor
,
E.
,
Tu
,
M.
,
Chia
,
D.
,
Spruce
,
R.
,
Batis
,
N.
,
Mehanna
,
H.
,
Wong
,
D. T. W.
, and
Huang
,
T. J.
(
2020
). “
Acoustofluidic salivary exosome isolation: A Liquid biopsy compatible approach for human papillomavirus-associated oropharyngeal cancer detection
,”
J. Mol. Diagn.
22
,
50
59
.
61.
Wu
,
M.
,
Chen
,
K.
,
Yang
,
S.
,
Wang
,
Z.
,
Huang
,
P. H.
,
Mai
,
J.
,
Li
,
Z. Y.
, and
Huang
,
T. J.
(
2018
). “
High-throughput cell focusing and separation via acoustofluidic tweezers
,”
Lab Chip
18
,
3003
3010
.
62.
Wu
,
J.
, and
Nyborg
,
W. L.
(
2008
). “
Ultrasound, cavitation bubbles and their interaction with cells
,”
Adv. Drug Deliv. Rev.
60
,
1103
1116
.
63.
Wu
,
M.
,
Ouyang
,
Y.
,
Wang
,
Z.
,
Zhang
,
R.
,
Huang
,
P. H.
,
Chen
,
C.
,
Li
,
H.
,
Li
,
P.
,
Quinn
,
D.
,
Dao
,
M.
,
Suresh
,
S.
,
Sadovsky
,
Y.
, and
Huang
,
T. J.
(
2017
). “
Isolation of exosomes from whole blood by integrating acoustics and microfluidics
,”
Proc. Natl. Acad. Sci. U.S.A.
114
,
10584
10589
.
64.
Yang
,
Y.
,
Wang
,
L.
,
Bell
,
P.
,
McMenamin
,
D.
,
He
,
Z.
,
White
,
J.
,
Yu
,
H.
,
Xu
,
C.
,
Morizono
,
H.
,
Musunuru
,
K.
,
Batshaw
,
M. L.
, and
Wilson
,
J. M.
(
2016
). “
A dual AAV system enables the Cas9-mediated correction of a metabolic liver disease in newborn mice
,”
Nat. Biotechnol.
34
,
334
338
.
65.
Yeo
,
L. Y.
,
Chang
,
H. C.
,
Chan
,
P. P.
, and
Friend
,
J. R.
(
2011
). “
Microfluidic devices for bioapplications
,”
Small
7
,
12
48
.
66.
Zhang
,
B.
,
Shi
,
Y.
,
Miyamoto
,
D.
,
Nakazawa
,
K.
, and
Miyake
,
T.
(
2019
). “
Nanostraw membrane stamping for direct delivery of molecules into adhesive cells
,”
Sci. Rep.
9
,
6806
.
67.
Zhao
,
S.
,
Borden
,
M.
,
Bloch
,
S. H.
,
Kruse
,
D.
,
Ferrara
,
K. W.
, and
Dayton
,
P. A.
(
2004
). “
Radiation-force assisted targeting facilitates ultrasonic molecular imaging
,”
Mol. Imaging
3
,
135
148
.
68.
Zhou
,
M.
,
Gao
,
D.
,
Yang
,
Z.
,
Zhou
,
C.
,
Tan
,
Y.
,
Wang
,
W.
, and
Jiang
,
Y.
(
2021
). “
Streaming-enhanced, chip-based biosensor with acoustically active, biomarker-functionalized micropillars: A case study of thrombin detection
,”
Talanta
222
,
121480
.
69.
Zhou
,
Y.
,
Shi
,
J.
,
Cui
,
J.
, and
Deng
,
C. X.
(
2008
). “
Effects of extracellular calcium on cell membrane resealing in sonoporation
,”
J. Control Rel.
126
,
34
43
.