FAST‐CRISPR: Fusogenic Association and Secured Transfection of CRISPR/Cas9 Ribonucleoproteins Using Lipid‐Silica Hybrid Nanoparticles for Therapeutic Genome Editing

ABSTRACT

Clinical translation of CRISPR/Cas9 therapeutics is challenged by inefficient cytosolic delivery and toxicity issues associated with viral vectors and nanoparticle‐based carriers. To overcome these concerns, herein we report a lipid‐silica hybrid nanoparticle platform for fusogenic association and secured transfection of CRISPR/Cas9 (FAST‐CRISPR), designed for rapid cytosolic delivery of CRISPR/Cas9 ribonucleoproteins, followed by efficient gene editing. Through direct fusion with the plasma membrane and bypassing conventional endocytic barriers, FAST‐CRISPR nanoparticles displayed superior intracellular delivery efficacy. Optimizing lipid compositions, we discovered that a 1:1 weight mixture of cationic DOTAP and ionizable DODMA lipids, combined with tailored large‐pore silica nanoparticles, enables enhanced loading capacity, rapid cytosolic dispersion, and significant nuclear transport of Cas9/gRNA complexes. FAST‐CRISPR nanoparticles efficiently delivered multiplex genome‐targeting ribonucleoproteins to induce targeted double‐strand DNA breaks, triggering apoptosis in cancer cells and significantly suppressing tumor growth in a mouse xenograft model without systemic toxicity. Our findings demonstrate the therapeutic efficacy and translational potential of FAST‐CRISPR nanoparticles as a safe and versatile non‐viral delivery platform for precision genome editing.

FAST‐CRISPR nanoparticles combine fusogenic lipids and tailored porous silica for rapid and direct cytosolic delivery of CRISPR/Cas9 ribonucleoproteins, bypassing endocytosis. This approach achieves highly efficient genome editing, selectively induces cancer cell apoptosis through multiplex DNA targeting, and effectively suppresses tumor growth in vivo, highlighting its translational potential as a safe, versatile, nonviral platform for precision genome editing.

1. Introduction

CRISPR/Cas9‐based gene editing technology represents a transformative advancement in medicine, enabling precise correction of disease‐causing mutations and offering unprecedented potential for treating diverse genetic disorders [1, 2, 3, 4]. Although viral vectors have been extensively utilized for CRISPR/Cas9 delivery due to their intrinsic high transduction efficiency, clinical translation is significantly hindered by safety concerns such as immunogenicity, insertional mutagenesis arising from random genomic integration [5]. Furthermore, prolonged Cas9 expression facilitated by viral vectors exacerbates the risk of off‐target gene editing and unintended DNA damage [6], emphasizing the critical need for the development of effective and safe non‐viral delivery alternatives for clinical translation.

Non‐viral delivery strategies typically involve CRISPR/Cas9 systems in forms of plasmid DNA, mRNA, or ribonucleoprotein (RNP) complexes. Plasmid DNA‐based methods, similar to viral vectors, carry heightened risks of off‐target effects due to prolonged Cas9 expression [7]. Delivery of Cas9 as mRNA mitigates genomic integration risks, but suffers from inherent instability, limited translation efficiency, and innate immunogenicity mediated by Toll‐like receptors (TLRs) [8, 9]. In contrast, intracellular delivery of Cas9 protein complexed with guide RNA (gRNA) as a ribonucleoprotein (RNP) form has emerged as a promising alternative, providing immediate editing activity without the translational constraints associated with mRNA. The transient nature and rapid degradation of RNP complexes substantially reduce off‐target gene editing, sustained DNA damage, and TLR‐mediated immune responses, collectively enhancing safety and efficacy [10, 11, 12, 13, 14].

Numerous non‐viral vehicles, including virus‐like particles, polymeric carriers, inorganic nanoparticles, and lipid nanoparticles (LNP), have been developed for CRISPR/Cas9 RNP delivery [15, 16, 17, 18, 19, 20, 21, 22, 23]. However, most of these delivery carriers depend heavily on endocytic uptake, subsequently requiring efficient endosomal escape to release the cargo into the cytosol, followed by nuclear localization where gene editing takes place. Endosomal escape constitutes a critical bottleneck, often leading to substantial degradation of the payloads [24, 25, 26, 27, 28, 29, 30]. Furthermore, lysosomal proteases released into the cytosol during endosomal escape can trigger NLRP3 inflammasome activation, inflammation, and potential cytotoxicity due to apoptotic pathways stimulation [31]. To address these challenges, nonviral delivery strategies capable of endocytosis‐independent cellular uptake have gained considerable attention. For instance, fusogenic liposomes, consisted of cationic lipids combining with membrane fusion‐favorable neutral helper phospholipids, can facilitate direct fusion with the plasma membrane, thereby bypassing endocytic pathways and associated endosomal degradation [32, 33, 34, 35, 36]. Membrane fusion‐derived direct cytosolic delivery strategy not only markedly increase the transfection efficacy but also reduces the risk of inflammation and cellular toxicity associated with lysosomal disruption.

In this study, we present a novel non‐viral delivery system, termed FAST‐CRISPR (Fusogenic Association and Secured Transfection of CRISPR/Cas9), designed for rapid and efficient cytosolic delivery of CRISPR/Cas9 RNP complexes through direct membrane fusion bypassing endosomal trafficking. Leveraging porous silica nanoparticles with substantially enlarged pore size (∼ 20 nm), the FAST‐CRISPR platform ensures robust loading and stability of RNP complexes (Figure 1). Porous silica nanoparticles are frequently used to deliver small‐molecule drugs because their porous structure provides a high surface‐to‐volume ratio and, together with their biocompatibility, supports drug loading without significant toxicity or immunogenicity [37, 38, 39, 40]. Conventional mesoporous silica nanoparticles exhibit limited uses in RNP delivery due to a small pore size, typically less than 10 nm, which is not adequate to sufficiently load the RNPs into the porous cavity [41, 42, 43, 44, 45]. Endosomal entrapment of the nanoparticles during intracellular uptake also causes insufficient cytosolic delivery efficiency. In order to increase the cytosolic delivery of RNP complexes, we hypothesized that encapsulation of the RNP‐loaded silica nanoparticles with optimized fusogenic lipids would significantly enhance delivery performance by directly mediating membrane fusion. Unlike previously reported lipid‐coated silica nanoparticles or conventional LNPs, which predominantly rely on endocytosis due to suboptimal lipid compositions [46, 47, 48], FAST‐CRISPR nanoparticles employ an optimized formulation of cationic and neutral fusion‐promoting lipids, as well as ionizable lipid. Particularly, compared to previous works, utilization of ionizable lipid is one of the key features, leading to efficient release of RNP payloads upon fusogenic intracellular uptake through transient charge‐conversion. This formulation of nanoparticles is eligible for fusogenic association and secured transfection of CRISPR/Cas9, which rapidly deliver the RNP complexes in cytosol bypassing endocytosis, resulting in significantly accelerated RNP delivery kinetics and improved gene editing efficiency. Design, optimization, and mechanistic underpinnings of FAST‐CRISPR nanoparticles were systematically characterized, demonstrating their superior intracellular delivery kinetics and enhanced gene‐editing efficiency. Furthermore, we highlight the therapeutic potential of FAST‐CRISPR nanoparticles by delivering multiplex genome‐targeting RNP complexes designed to induce targeted double‐strand DNA breaks and apoptosis specifically in cancer cells [49, 50]. Through comprehensive in vitro validation and in vivo tumor suppression studies, our findings establish FAST‐CRISPR nanoparticles as a versatile and clinically promising genome‐editing delivery platform, offering significant advantages in precision, safety, and translational potential compared to existing delivery modalities.

FIGURE 1.

Schematic illustrations depicting the fusogenic association and secured transfection of CRISPR/Cas9 (FAST‐CRISPR) using lipid/silica hybrid nanoparticles. CRISPR ribonucleoprotein (RNP) is loaded into porous cavity of silica nanoparticles and further encapsulated with lipid layer, leading to fusogenic association with plasma membrane of the cell, followed by direct cytosolic delivery for secured transfection of RNP payloads. Optimal ratio of DOTAP (cationic), DODMA (ionizable), DOPE (membrane‐fusion favorable), and DSPE‐PEG (stabilizing) lipid components is responsible for fusogenic association of nanoparticles bypassing endosomal trafficking. In contrast, typical formulations of nanoparticles including lipid nanoparticles (LNPs), liposomes, and bare silica nanoparticles are entrapped in endosomes followed by lysosomal or autophagic degradation that subsequently limits the cytosolic delivery efficacy of therapeutic payloads such as CRISPR RNP or mRNA.

2. Results and Discussion

2.1. Tailored Design and Preparation of Lipid‐Silica Hybrid Nanoparticles for Efficient RNP Delivery

To demonstrate the utility of fusogenic lipids for intracellular delivery of CRISPR/Cas9 RNP, we initially examined fusogenic liposomes loaded with RNP or doxorubicin, a model small‐molecule drug. Although both payloads exhibited higher cellular association when delivered using fusogenic liposomes, sufficient translocation into the cytosol was only observed in doxorubicin, while RNPs remained predominantly trapped at the plasma membrane or endosomes (Figure S1). We reasoned that a strong electrostatic interaction between RNP and cationic lipid in the fusogenic liposome would hamper cytosolic release of RNP, while weakly bound doxorubicin can be readily released into cytosol after membrane fusion [51]. Thus, we hypothesized that distinct host carriers along with fusogenic lipid coating capable of securing loading and facile cytosolic release of RNP payloads are required as a part of delivery vehicles.

Given that the hydrodynamic diameter of RNP complexes is approximately 10 nm [52, 53, 54], we strategically designed porous silica nanoparticles featuring substantially larger pore sizes (∼20 nm) to efficiently accommodate RNPs (Figure 2a). The porous silica nanoparticles were synthesized using a modified Stöber reaction, employing sodium salicylate and cetyltrimethylammonium chloride as templating agents, generating large‐pores, triethanolamine as a catalyst, and tetraethyl orthosilicate as a silica precursor. Brunauer‐Emmett‐Teller (BET) analysis confirmed an average pore size of approximately 22 nm, a total surface area of 679.7 m² g⁻¹, and a Barrett‐Joyner‐Halenda (BJH) pore volume of 3.66 cm³ g⁻¹, which were higher than that of porous silica nanoparticles synthesized in longer reaction time (486 m² g⁻¹ and 1.92 cm³ g⁻¹ respectively) [55, 56, 57], indicating the suitability of porous nanoparticles for efficient loading of the RNP payload.

FIGURE 2.

Characterization of lipid/silica hybrid nanoparticle. (a) Schematic illustrations showing RNP loading and lipid encapsulation in the process preparing FAST‐CRISPR nanoparticles. Porous cavity (diameter of each pore (d _pore): ∼ 20 nm) of the silica nanoparticles is readily occupied with RNPs, and further coated with mixed ratio of lipids. (b) Transmission electron microscopy (TEM) images of (i) bare porous silica, (ii) RNP‐loaded silica, (iii) RNP‐loaded and lipid‐coated silica, and (iv) RNP‐loaded liposome (no silica core) nanoparticles, respectively. Note that lipid contents are negatively stained with UranyLess EM stain for clear distinction. Scale bar: 100 nm. Inset in (iv): Schematic of RNP‐loaded liposome. (c) Hydrodynamic size and (d) zeta potential of each formulation of nanoparticles (n = 3). (e) RNP loading amount per 1 µg of silica nanoparticle or equivalent mass of lipid used in encapsulation of silica nanoparticles (n = 3). (f) RNP loading amount and loading efficiency depending on the amount of RNP mixed (input RNP) with 1 µg of silica nanoparticle (n = 3). Data represented as mean ± standard deviation in (c–f).

Following nanoparticle synthesis, RNP complexes were effectively loaded into the silica nanopores through electrostatic interactions facilitated by inherent charge differences between the mixed‐charge RNPs and the negatively charged (hydroxyl‐terminated) silica surface. Subsequently, the RNP‐loaded silica nanoparticles were encapsulated with fusogenic lipids composed of DOTAP (cationic lipid), DODMA (ionizable lipid), DOPE (fusogenic neutral helper lipid), and DSPE‐PEG (stabilizing lipid) in weight ratio of 22.25:22.25:44.5:11 (Figures 1 and 2a). The ionizable lipid DODMA is structurally similar to DOTAP, thus chosen for its ability to form a stable, charge‐reversible lipid bilayer coating on silica nanoparticles, enabling efficient cargo release upon fusion with the cellular membrane [58]. The resulting lipid‐silica hybrid nanoparticles have a core‐shell structure comprising an RNP‐loaded porous silica nanoparticle in the core and fusogenic lipid shell engineered for direct cytosolic delivery, thus circumventing the typical bottleneck of endosomal entrapment and degradation.

Transmission electron microscopy (TEM) clearly visualized the well‐defined porous structure to ensure secured loading of RNP and confirmed successful lipid encapsulation around the nanoparticle (Figure 2b). TEM images explicitly differentiated lipid‐silica hybrid nanoparticles from bare silica, RNP‐loaded silica, and liposome alone without a silica core. Hydrodynamic size measurement of each nanoparticle formulation revealed uniform and consistent size distributions (Figure 2c), consistent with TEM observation. Zeta potential of nanoparticles further supported the sequential steps of RNP loading and lipid encapsulation, demonstrating substantial shifts toward neutral surface charges after lipid coating—closely comparable to that of liposome alone—indicating successful lipid encapsulation (Figure 2d).

Notably, we observed no significant difference in RNP loading amounts before and after lipid encapsulation, confirming negligible leakage during the lipid coating process (Figure 2e). In addition, the silica nanoparticles demonstrated over two‐fold higher RNP loading capacity compared to liposomes lacking a silica core, emphasizing the substantial benefit of the porous structure in highly retaining RNP payloads. We further investigated the influence of varying initial RNP concentrations on loading outcomes, indicating that increasing initial RNP concentrations for mixing with porous silica nanoparticles elevated total loaded RNP quantity but concurrently reduced overall loading efficiency (Figure 2f). This highlights a practical trade‐off that careful selection of initial RNP concentrations is critical for optimizing therapeutic efficacy per nanoparticle while maintaining cost‐effectiveness, tailored to specific therapeutic applications.

2.2. Fusogenic Association and Secured Transfection of CRISPR/Cas9 (FAST‐CRISPR)

We then assessed the cellular uptake and intracellular localization of fusogenic lipid‐coated porous silica nanoparticles capable of direct membrane fusion‐derived cytosolic delivery of RNP. Confocal laser scanning microscopy (CLSM) imaging revealed significant intracellular uptake of RNP in 2 h post‐treatment with lipid‐silica hybrid nanoparticles (Figure 3a). Particularly, the mixed lipid combination of DOTAP and DODMA (1:1 weight ratio) was highly beneficial for both intracellular uptake and nuclear localization of RNPs (Figure 3b,c). In contrast, liposomes without the silica core showed noticeable binding to the plasma membrane with limited cytosolic release and negligible nuclear accumulation of RNPs. This result suggested that strong electrostatic interactions between lipids and RNPs might limit their dissociation and subsequent intracellular distribution even after membrane binding and fusion. However, the silica nanoparticles facilitated efficient cytosolic release and nuclear localization of RNPs, as RNP‐loaded silica nanoparticles exhibited less electrostatic interactions with plasma membrane after cellular uptake.

FIGURE 3.

Intracellular delivery of FAST‐CRISPR nanoparticles. (a) Confocal laser scanning microscopy (CLSM) images of HCT‐116 cells obtained at 2 h post‐treatment with nanoparticles or free RNP as indicated. Note that the ratio of DOTAP:DODMA is 1:1 in DOTAP+DODMA groups in (iv) and (viii). Atto550 and DiD molecules are labeled to gRNA (red) and lipid (green), respectively. Scale bar: 20 µm. (b and c) Quantification of gRNA based on fluorescence signal per cell showing the localization in (b) intracellular region (both in cytosol and nucleus) or (c) nucleus (n = 3). (d and e) Indel frequency in EMX1 genomic locus of HCT‐116 cells obtained at 3 days post‐treatment with nanoparticles or free RNP as indicated (n = 3). Inset: T7 endonuclease I assay gel images. (f) CLSM images showing cellular distribution of atto550‐labeled gRNA representing RNP delivery using 50:50 ratio of DOTAP:DODMA encapsulated silica nanoparticles (FAST‐CRISPR) or LNP at different incubation time as indicated. Scale bar: 20 µm. (g) Corresponding fluorescence intensity ratio of gRNA in nucleus to cytosol after RNP delivery using FAST‐CRISPR or LNP at different incubation time. Data represented as mean ± standard deviation and statistically analyzed via one‐way ANOVA and Tukey's post hoc test; ns, not significant; ^* p < 0.05; ^** p < 0.01; ^**** p < 0.0001.

To further elucidate the surface charge effect of silica nanoparticles, we further demonstrated the RNP delivery with amine‐functionalized (positively charged) silica nanoparticles after encapsulation with fusogenic lipids. Amine‐functionalized silica nanoparticles showed markedly reduced cytosolic RNP delivery (Figure S2), presumably due to strong binding affinity to the inner membrane of the cell, even after cell uptake, thereby restricting cytosolic distribution of the payloads. Moreover, we speculate that a positively charged surface possibly interferes with stable lipid encapsulation and intensifies nonspecific interactions with the plasma membrane, followed by hindering effective RNP release into the cytosol. Overall, silica nanoparticles not only provide sufficient volume for high loading capacity of RNP payloads but also lead to surface charge‐dependent facile cytosolic distribution upon intracellular uptake.

Since we used a fusogenic neutral helper lipid (DOPE) as a part of lipid encapsulants, both DOTAP‐coated and DODMA‐coated nanoparticles (formulation (vi) and (vii), respectively) also showed promising endocytic delivery of RNP payloads, owing to membrane‐fusion derived intracellular uptake (Figure 3a–c). However, it should be also noted that both cellular uptake and nuclear localization were significantly enhanced by using mixed lipid combination of DOTAP and DODMA rather than single use of cationic or ionizable lipid for silica nanoparticle encapsulation. Even though ionizable lipids are developed to overcome the limitations of permanent cationic lipids, such as strong charge interaction with nucleic acids as well as low encapsulation and delivery efficiency [51, 59, 60, 61], permanently cationic lipid at physiological pH is still essential for higher cellular uptake. In this regard, cationic lipid such as DOTAP is frequently added to LNP formation in addition to ionizable lipid to enhance delivery efficiency [54, 62]. However, those LNPs still suffer from endosomal degradation and limited cytosolic delivery. In the present study, applying the silica nanoparticles in the core of lipid layer, we achieved substantially enhanced intracellular delivery of RNPs through direct membrane fusion and efficient cytosolic release, resulting in superior transport to nucleus as well.

Since the nuclear localization signal (NLS) tags on Cas9 protein facilitate the entry into the nucleus, cytosolic delivery of RNPs is crucial for efficient CRISPR/Cas9 genome editing in the cell. Therefore, we further validated gene editing efficiency upon RNP delivery, which showed markedly higher indel frequency with DOTAP‐DODMA mixed lipid‐coated silica nanoparticle compared to other formulations (Figure 3d). Although noticeable nuclear localization of RNP was also found with either DOTAP or DODMA‐coated silica nanoparticles (Figure 3a–c), negligible gene editing was detected in those groups, possibly due to insufficient delivery amount less than the threshold for effective gene editing. To reveal the optimized lipid ratio of cationic and ionizable lipids, gene editing efficacy upon RNP delivery with lipid‐silica hybrid nanoparticle varying ratio of DOTAP and DODMA was also quantitatively evaluated. The result showed the highest gene editing efficiency when 50:50 ratio of DOTAP and DODMA lipids was used (Figure 3e). Moreover, we found that DODMA has proper pK_a value (∼6.59) for superior membrane binding and fusogenic cellular characteristics to be used with DOTAP, as a lower pK_a valued ionizable lipid (e.g., DODAP, pK_a∼5.59) might have less chance to be positively charged in the cellular uptake condition (Figure S3). Consequently, the optimized formulation for fusogenic association and secured transfection of CRISPR/Cas9 (FAST‐CRISPR) comprises silica nanoparticles loaded with RNPs and coated with 50:50 mixture of DOTAP and DODMA lipids.

To validate the advantages of FAST‐CRISPR nanoparticles over endocytosis‐dependent lipid nanoparticles (LNPs), we compared FAST‐CRISPR with a representative LNP formulation using the well‐established ionizable lipid C12‐200 at an optimized molar ratio of C12‐200:DOPE:cholesterol:DMG‐PEG (35:16:46.5:2.5), with an additional 10% DOTAP. This LNP is widely used as a gold standard for screening and has demonstrated robust gene‐delivery activity in vitro and in vivo including in rodents and nonhuman primates, particularly for liver delivery [54, 63, 64, 65]. Although conventional LNPs efficiently encapsulate and deliver nucleic acids (or RNPs), their reliance on endosomal escape mechanisms exposes delivered cargo to risks of lysosomal degradation and delayed nuclear localization [66, 67]. In contrast, FAST‐CRISPR nanoparticles demonstrated significantly higher nuclear to cytosol ratio of RNP delivery (Figure 3f,g; Figure S4). Quantitative tracking over time indicated more than two‐fold increase in nuclear to cytosol ratio of RNP accumulation with FAST‐CRISPR nanoparticles compared to conventional LNPs (Figure 3g). The rapid nuclear delivery facilitated by FAST‐CRISPR was attributable to its distinct membrane fusion mechanism, entirely bypassing endosomal pathways. Thus, the NLS‐containing Cas9 protein maintains integrity and functionality, avoiding degradation encountered during prolonged endosomal transit. Furthermore, time‐course tracking in a short‐period post‐treatment corroborated that nuclear accumulation of RNPs occurred significantly faster with FAST‐CRISPR nanoparticles than LNPs (Figure S5) whereas at later time points the nuclear accumulation reached similar levels (Figure S4), possibly owing to direct membrane fusion in contrary to sequential processes of endocytosis, endosome acidification and endosomal escape via endosomal membrane disruption. The accelerated delivery is essential not only for efficient genome editing but also to reduce potential off‐target effects resulting from prolonged exposure to intracellular environments.

Collectively, these results highlight that the lipid composition critically influences the membrane binding of carriers and intracellular fate of delivered RNPs. By strategically balancing cationic DOTAP and ionizable DODMA lipids within a fusogenic lipid‐coated porous silica nanoparticle platform, we achieved rapid, secure, and efficient cytosolic delivery of CRISPR/Cas9 RNPs. The resulting FAST‐CRISPR nanoparticles circumvent classical endocytic bottlenecks, offering substantial improvements in nuclear localization of RNPs and gene editing efficiency compared to conventional LNP systems. Given these advantages, the FAST‐CRISPR nanoparticle system presents strong translational potential for precise and rapid genome editing, particularly suited for therapeutic applications requiring high efficiency.

2.3. Intracellular Delivery Mechanism of FAST‐CRISPR

To elucidate the intracellular delivery mechanism of FAST‐CRISPR nanoparticles, we further conducted time‐lapse CLSM imaging at single cell level. Upon exposure to FAST‐CRISPR nanoparticles, rapid cytosolic dispersion of RNPs was observed, notably absent of early endosomal accumulation, indicating an endocytosis‐independent delivery pathway (Figure 4a). Initially, distinct fluorescent puncta indicative of nanoparticles were associated with the plasma membrane, which rapidly dissipated, allowing RNPs to uniformly distribute throughout the cytosol and subsequently enter the nucleus. In contrast, LNPs initially displayed pronounced endosomal accumulation followed by gradual cytosolic dispersion only at later stages as typical evidence of endocytosis (Figure S5b). Interestingly, FAST‐CRISPR nanoparticles exhibited an opposite temporal profile: rapid initial cytosolic dispersion was observed, with only minor accumulation in intracellular vesicles becoming apparent at later stages (Figure S5a). Despite predominantly fusogenic uptake, such late‐stage presence of punctate RNPs likely resulted from secondary autophagic sequestration of cytosolically delivered cargo or minor unintended cellular uptake via endocytosis. Additional CLSM imaging performed across different focal planes comprehensively evidenced extensive intracellular delivery of both Cas9 protein and gRNA, clearly supporting the efficient cytosolic delivery and nuclear localization facilitated by FAST‐CRISPR nanoparticles (Figure 4b,c).

FIGURE 4.

Rapid cytosolic delivery of RNPs using FAST‐CRISPR. (a) Time‐lapse CLSM images of fluorescent RNP trafficking at the single‐cell level. Note that the membrane‐bound nanoparticles (lined circles) are directly fused with cell membrane, and then RNP and DiD (labeled for lipid) are quickly distributed in cytosol and cell membrane, respectively, as fluorescence puncta in the same positions (dotted circles) were substantially reduced in 15 min. Scale bar: 10 µm. (b) Reconstructed 3D image at a focal plane showing both cytosolic (yellow arrows) and nuclear (white arrows) delivery of RNPs with membrane‐bound nanoparticles (green arrows) at 2 h post‐treatment with FAST‐CRISPR nanoparticles. (c) CLSM images of a HCT‐116 cell at different focal planes (z‐stacked) obtained at 2 h post‐treatment with FAST‐CRISPR nanoparticles loaded with RNPs (dCas9‐GFP fusion protein and atto550‐labeled gRNA). (d and e) Intracellular uptake of FAST‐CRISPR nanoparticles upon pre‐treatment (1 h) with endocytosis inhibitors: Genistein (GEN), caveolae‐mediated endocytosis inhibitor; Chloropromazine (CPZ), clathrin‐mediated endocytosis inhibitor; Amiloride (AMI), macropinocytosis inhibitor. (d) CLSM images of atto550‐labeled gRNA (red) obtained at 1.5 h post‐treatment with FAST‐CRISPR nanoparticles, and (e) corresponding quantification of gRNA localized in nucleus (i.e., white dotted regions in (d)) (n = 3). Scale bar: 20 µm. (f–i) Comparison of RNP delivery using FAST‐CRISPR with membrane fusion favorable helper lipid DOPE against endocytic helper lipid DOPC. (f) CLSM images showing higher cytosolic delivery of gRNA (red) with DOPE‐based FAST‐CRISPR nanoparticles rather than DOPC‐based one at 2 h post‐treatment. Scale bar: 20 µm. (g) Corresponding gRNA quantification localized in nucleus (n = 3). (h) T7 endonuclease I assay gel image and (i) corresponding indel frequency in EMX1 genomic locus of HCT‐116 cells 3 days post‐treatment (n = 3). Data represented as mean ± standard deviation and statistically analyzed via one‐way ANOVA and Tukey's post hoc test for (e) and two‐tailed unpaired Student's t‐test for (g and i); ns, not significant; ^*** p < 0.001.

To further verify the independence of the FAST‐CRISPR delivery mechanism from endocytosis, we performed experiments involving pharmacological inhibition of major endocytic pathways. Cells were preincubated with inhibitors targeting clathrin‐mediated (chlorpromazine), caveolae‐mediated (genistein), and macropinocytosis‐dependent (amiloride) endocytosis pathways. Notably, none of these inhibitors significantly affected the intracellular delivery efficiency of RNPs using FAST‐CRISPR nanoparticles, as evidenced by robust cytosolic localization unaffected by inhibitor treatment (Figure 4d,e). In contrast, intracellular delivery of RNPs by the LNPs was substantially inhibited by macropinocytosis‐dependent inhibitor amiloride (Figure S6). This clearly demonstrates that the intracellular delivery of FAST‐CRISPR nanoparticle is independent of the conventional endocytic mechanism.

In addition, to assess the critical role of lipid‐mediated membrane fusion in FAST‐CRISPR nanoparticles, we prepared control formulations replacing the DOPE (fusion‐favorable helper lipid) with DOPC (endocytic helper lipid) while other lipid contents were maintained. Significantly lower delivery efficacy of RNPs in the nucleus was observed with DOPC‐based nanoparticles (Figure 4f,g), which also failed to induce noticeable gene editing efficiency, correspondingly (Figure 4h,i). These results unequivocally support that the enhanced intracellular delivery of RNPs by FAST‐CRISPR nanoparticles specifically arises from fusogenic lipid‐mediated direct membrane fusion rather than endocytic uptake. Taken together, these findings convincingly demonstrate that FAST‐CRISPR nanoparticles efficiently and rapidly deliver the RNP payloads via direct plasma membrane fusion, bypassing traditional endocytic bottlenecks. This mechanism not only enhances delivery kinetics but also minimizes endosomal degradation risks, highlighting the considerable potential of FAST‐CRISPR for efficient genome‐editing applications.

2.4. Therapeutic Efficacy of FAST‐CRISPR Delivering Cancer‐Specific Indel Attacker In Vitro

Having established FAST‐CRISPR nanoparticles as an efficient platform enabling rapid cytosolic delivery of RNPs, we next sought to evaluate their therapeutic potential in cancer cells. For this purpose, we employed a DNA‐damaging CRISPR/Cas9 system termed cancer‐specific indel attacker (CINDELA) [49], designed to simultaneously induce extensive genomic double‐strand breaks (DSBs) at multiple genomic loci (50 distinct target sites in this study), triggering apoptotic cell death (Figure 5a). Unlike conventional single‐gene editing approach, multiplexed targeting of numerous genomic loci in CINDELA strategy amplifies the extent of DNA damage, effectively activating apoptotic pathways. Although the CINDELA RNP used in this study does not induce DSBs in a cancer‐specific manner, the multiplexed genome targeting strategy also holds promise for personalized cancer therapy by selectively inducing lethal DSBs at cancer‐specific or patient‐specific mutated genomic loci, thereby ensuring targeted therapeutics outcomes with minimal off‐target effects [68].

FIGURE 5.

Therapeutic anti‐cancer efficacy of DNA‐damaging CRISPR/Cas9 RNP delivery in vitro. (a) Schematic illustrations of cancer cell death induced by double‐strand breaks at multiple target genomic loci generated from delivery of target DNA‐damaging RNPs using FAST‐CRISPR nanoparticles. (b) CLSM images showing DNA damages (white) in the nucleus (blue) of HCT‐116 cells through immunofluorescence staining with γH2AX at 24 h post‐treatment with free RNP or FAST‐CRISPR nanoparticles loaded with either target RNP (T@FAST‐CRISPR) or non‐target RNP (NT@FAST‐CRISPR). Scale bar: 20 µm (c) Corresponding quantification of γH2AX fluorescence signals in nucleus (n = 3). (d) LIVE/DEAD staining images of HCT‐116 cells at 3 days post‐treatment as indicated: Nucleus (blue), live cells (green), and dead cells (red). Scale bar: 50 µm. (e) Cell viability measured using CCK‐8 assay at 3 days post‐treatment with either non‐target or target RNP loaded in different formulations of nanoparticles as indicated (n = 3). Data represented as mean ± standard deviation and statistically analyzed via one‐way ANOVA for (c) and two‐way ANOVA for (e) and Tukey's post hoc test for both; ns, not significant; ^**, p < 0.01; ^*** p < 0.001; ^****, p < 0.0001.

To validate CINDELA‐mediated genomic damage, we assessed the induction of genomic DNA damage through immunofluorescence staining of γH2AX, a well‐established marker of DSBs, upon delivery of RNPs (Figure 5b,c). Target RNP was designed to induce the DSBs at 50 genomic loci in cancer cell DNA, thus target RNP‐loaded FAST‐CRISPR nanoparticles (T@FAST‐CRISPR) induced substantial γH2AX formation, indicative of widespread genomic damage. However, delivery of nontarget RNP (designed to generate a single DSB at the EMX1 locus) using FAST‐CRISPR nanoparticles (NT@FAST‐CRISPR) showed minimal γH2AX activation, comparable to control groups. Correspondingly, live/dead staining and viability assays further confirmed the therapeutic effect of CINDELA, displaying substantial cell death exclusively in cancer cells treated with T@FAST‐CRISPR nanoparticles among various formulations of RNP‐loaded nanoparticles (Figure 5d). As nontarget RNP showed negligible DNA damage and cytotoxicity, the result also implies nontoxic nature of FAST‐CRISPR nanoparticles as a RNP carrier as well. In addition, 50:50 mixture of DOTAP and DODMA lipids showed highest cell death efficacy (Figure S7), in coincidence with indel frequency of gene editing RNP delivery (Figure 3e).

Further highlighting the advantages of FAST‐CRISPR, we compared the therapeutic performance and intrinsic toxicity with LNPs. FAST‐CRISPR nanoparticles without RNP payloads displayed negligible cytotoxicity themselves, while equivalent LNPs showed severe toxicity at higher dosage (Figure S8). Although FAST‐CRISPR also showed slightly reduced cell viability at higher dosage, we confirmed that FAST‐CRISPR nanoparticles are potentially a safer and more efficient delivery platform. It should be also noted that both FAST‐CRISPR and LNPs showed comparable cell viability while FAST‐CRISPR enabled rapid and efficient intracellular delivery of RNPs. This result supports that the total amount of nuclear‐localized RNPs was presumably equivalent at the end, featuring rapid cellular uptake and secured transport of RNP payloads, corresponding to the above findings (Figure S5). These characteristics were attributed from fusogenic delivery mechanism, bypassing endocytosis and subsequent lysosomal rupture—events strongly associated with inflammatory responses and cell death commonly observed with conventional LNPs [69, 70, 71]. Because endosomal escape and ensuing inflammatory cascades drive pro‐inflammatory cytokine secretion, FAST‐CRISPR is expected to elicit a lower in vivo inflammatory response, thereby supporting its safer translational profile [72, 73, 74]. Despite the beneficial features of FAST‐CRISPR for rapid and efficient intracellular delivery, extensive studies on safety profiles and cytotoxicity upon long‐term exposure should be further investigated in depth.

We also evaluated the morphological characteristics of silica nanoparticles on RNP delivery and therapeutic performance. Compared to smaller‐sized porous silica nanoparticles (size: ∼100 nm; pore diameter: ∼20 nm) coated with identical lipid composition, our FAST‐CRISPR nanoparticles (size: ∼360 nm; pore diameter: ∼22 nm) demonstrated superior anti‐cancer efficacy (Figure S9). Although smaller‐sized nanoparticles are supposed to be advantageous in cellular uptake [75, 76], higher therapeutic outcome is likely attributable to the larger surface area (679.7 m² g⁻¹) of the porous silica nanoparticles used in FAST‐CRISPR relative to smaller‐sized silica nanoparticles (488.0 m² g⁻¹), facilitating increased loading capacity in a single nanoparticle and more efficient intracellular release of RNPs. Moreover, incomplete or unstable encapsulation of smaller‐sized silica nanoparticles with the lipid layer, presumably due to steric hindrance and higher curvature of lipid layer on the surface, might reduce the fusogenic cellular uptake [77, 78, 79]. Collectively, our findings emphasize the FAST‐CRISPR nanoparticles as a tailor‐designed, highly potent, efficient, and safe delivery platform for multiplex genome‐targeting CRISPR/Cas9 RNPs, strongly supporting their potential as a versatile therapeutic approach for precision oncology applications.

2.5. In Vivo Therapeutic Efficacy of FAST‐CRISPR Nanoparticles for Tumor Suppression

To assess the translational therapeutic potential of FAST‐CRISPR nanoparticles, we evaluated their anti‐cancer efficacy through CINDELA RNP delivery in vivo using HCT‐116 xenograft tumor model established in immunodeficient nude mice. When tumor volumes reached approximately 80 mm³, mice received daily intratumoral injections of either T@FAST‐CRISPR, NT@FAST‐CRISPR nanoparticles, or phosphate‐buffered saline (PBS) alone as a control. Treatments were continued for 12 consecutive days, and both body weights and tumor volumes were monitored throughout the study period.

Upon treatment with T@FAST‐CRISPR nanoparticles, significant suppression of tumor growth was observed while other groups exhibited continuous tumor progression (Figure 6a). The potent anti‐cancer effect was further confirmed at day 14 through direct assessment of harvested tumor weight and size (Figure 6b,c). To verify the mechanism underlying the observed tumor suppression, tumor tissues were further analyzed by immunohistochemistry staining with γH2AX for DNA DSB, Caspase, and TUNEL assay for apoptotic cell death (Figure 6d; Figure S10a). The T@FAST‐CRISPR nanoparticles‐treated tumors exhibited significantly elevated levels of staining for γH2AX (Figure 6e), TUNEL (Figure 6f), and Caspase (Figure S10b), confirming successful induction of extensive genomic DNA damage followed by apoptosis within treated tumor tissues. These results aligned with the in vitro findings and support a robust mechanism of action in which rapid and efficient delivery of CINDELA RNPs by FAST‐CRISPR nanoparticles drives targeted genomic damage and apoptotic cell death. Importantly, daily administration of FAST‐CRISPR nanoparticles did not produce observable systemic toxicity, as indicated by stable body weights in all treated mice throughout the treatment duration (Figure S11). This highlights the inherent biocompatibility and minimal toxicity of the FAST‐CRISPR nanoparticle formulation, which was even more favorable than that of LNPs, likely resulting from its endocytosis‐independent, membrane fusion‐based delivery pathway that circumvents inflammatory and cytotoxic responses associated with endosomal disruption.

FIGURE 6.

Suppressing tumor growth with DNA‐damaging RNP delivery in vivo xenograft model. (a) Tumor volume tracking upon daily intratumoral injections of PBS or FAST‐CRISPR nanoparticles loaded with either target RNP (T@FAST‐CRISPR) or non‐target RNP (NT@FAST‐CRISPR). (n = 6 for PBS and n = 5 for other groups). (b) Harvested tumor weight and (c) photographs obtained at 14 days post‐treatment with PBS, T@FAST‐CRISPR (denoted as T), and NT@FAST‐CRISPR (denoted as NT). Scale bar: 1 cm. Note that FAST‐CRISPR nanoparticles contain DiD dye to visualize successful intratumoral injection, thus the tumor tissues appear to be greenish color. (d) Immunohistochemical staining of γH2AX (DNA damage marker) and TUNEL (apoptosis marker) on hematoxylin‐stained tumor tissues: Nucleus (blue), γH2AX or TUNEL (brown). Inset: Whole tissue section images of tumors, where 1 mm x 1 mm white box indicating location of the magnified images. Scale bar: 200 µm. (e and f) Staining positive area in whole area of tumor tissue sections. Data represented as mean ± standard deviation and statistically analyzed via two‐way ANOVA for (a) and one‐way ANOVA for (b, e, f) and Tukey's post hoc test for both; ns, not significant; ^* p < 0.05; ^** p < 0.01; ^**** p < 0.0001.

While intratumoral administration was chosen here to simplify the in vivo demonstration and to isolate on‐mechanism activity of FAST‐CRISPR, this route by itself does not adequately address deep‐seated or metastatic tumors. To map a translational path toward systemic application, we outline design strategies that are compatible with the current platform. First, the lipid shell is readily amenable to ligand integration, including aptamers, peptides, and antibodies to enable active tumor cell targeting and homing to both primary and metastatic lesions [62, 80, 81, 82, 83]. Second, pharmacokinetic optimization (e.g., tuning PEG density and chain length, cholesterol content, and tail saturation) can prolong circulation and reduce nonspecific interactions while preserving delivery competence. Third, conditional fusogenicity can be engineered to bias activity toward the acidic tumor microenvironment by adjusting the fraction of ionizable lipids in the shell and/or incorporating pH‐labile or charge‐reversal chemistries, thereby masking fusogenicity in blood and unmasking it in tumors [84, 85, 86, 87, 88]. Finally, tissue‐penetration parameters (hydrodynamic size in the ∼70–120 nm range, moderate surface charge, and matrix‐penetrating motifs) can be leveraged to improve intratumoral distribution beyond passive extravasation [89, 90, 91]. These considerations, together with the demonstrated potency and local safety of FAST‐CRISPR in vivo, support the feasibility of systemic, clinically relevant deployment. Overall, our study establishes the FAST‐CRISPR nanoparticles as a powerful and translationally promising platform for CRISPR‐based precision nanomedicine, providing rapid, efficient, and safe delivery of therapeutic genome editing agents in vivo, while offering strategic engineering routes to systemic application.

3. Conclusions

In this study, we developed a fusogenic lipid‐silica hybrid nanoparticle (named FAST‐CRISPR) for efficient cytosolic delivery of CRISPR/Cas9 RNPs, overcoming critical limitations associated with conventional viral vectors or LNP‐based delivery systems. By carefully optimizing the lipid composition—specifically a balanced mixture of cationic DOTAP and ionizable DODMA lipids (1:1 weight ratio)—alongside tailored large‐pore silica nanoparticles, we achieved direct cytosolic delivery via plasma membrane fusion, bypassing endosomal entrapment and significantly enhancing intracellular and nuclear localization of therapeutic RNPs. The membrane fusion‐derived delivery mechanism was confirmed through comprehensive studies, clearly demonstrating independence from traditional endocytic uptake pathways and providing improved safety and kinetics compared to typical LNPs. Using FAST‐CRISPR nanoparticles, we successfully delivered a multiplex DNA‐damaging CRISPR/Cas9 system to induce extensive genomic DNA damage and apoptosis in cancer cells, both in vitro and in vivo. In a xenograft tumor mouse model, FAST‐CRISPR nanoparticles significantly suppressed tumor growth without observable systemic toxicity, highlighting their therapeutic efficacy and biocompatibility. The versatility of FAST‐CRISPR technology was further demonstrated by its increased loading capacity, allowing enhanced therapeutic efficacy at lower nanoparticle doses as well. Overall, our findings establish FAST‐CRISPR nanoparticles as a potent and safe nonviral CRISPR/Cas9 delivery platform, with translational potential for precision oncology and personalized nanomedicine. Further developments, such as incorporating targeting ligands (e.g., aptamers and peptides) or exploiting the acidic tumor microenvironment through the pH‐dependent charge conversion of the ionizable lipid DODMA, could enable systemic delivery and facilitate clinical translation to deep‐seated and metastatic cancers. These strategies will broaden the therapeutic applications of FAST‐CRISPR technology as a universal CRISPR delivery platform not limited to RNP but also many engineered versions of CRISPR systems, including base and prime editors in forms mRNA with gRNA.

4. Experimental Section

4.1. Synthesis of Porous Silica Nanoparticles

In a glass vial, triethanolamine (TEA, 35 mg, Sigma‐Aldrich, 90279) and deionized water (12.5 mL) were mixed under 1000 rpm magnetic stirring at 80°C for 30 min. Then, sodium salicylate (NaSal, 84 mg, Millipore, 106601) and 25% (w/w) cetyltrimethylammonium chloride solution (CTAC, 520 µL, Sigma‐Aldrich, 292737) were added to the above solution and stirred for 1 h. After that, a mixture of tetraethyl orthosilicate (TEOS, 2.25 mL, Sigma‐Aldrich, 86578) and ethanol (0.25 mL, Millipore, 100983) was added dropwise and stirred for 2 h with 1000 rpm and 80°C throughout the procedure. The synthesized nanoparticles were collected and washed with ammonium nitrate (6 g L⁻¹, Sigma‐Aldrich, 221244) in ethanol for 4 times, 1% (w/w) hydrochloric acid (Junsei Chemical, 20 010–1250) in ethanol for 2 times and pure ethanol for 2 times by using 21 000 rcf centrifugation for 10 min and sonication in an ultrasonic bath (40 kHz, 50 W, SH‐501, Saehan Ultrasonic, Korea) for 15 min to remove unreacted reactants and surfactants. For each wash, the volume of the wash solution was equal to the synthesis reaction volume. The washed silica nanoparticle was completely dried at 80°C under vacuum and weighed.

4.2. Preparation of FAST‐CRISPR Nanoparticles

Ribonucleoprotein (RNP) complex of CRISPR/Cas9 and gRNA was formed by mixing equimolar Cas9 protein with gRNA in PBS and 10 min incubation at room temperature. Then porous silica nanoparticle (2 mg mL⁻¹) dispersed in PBS was added to RNP complex and gently shaken at 500 rpm in 4°C for 15 min using Thermomixer C (Eppendorf, Germany). Directly after, lipid mixture for fusogenic liposome (20 mg mL⁻¹) in ethanol was mixed with RNP‐loaded silica nanoparticle by rapid pipetting and incubated for 10 min at room temperature. The free lipid and RNP in the supernatant were washed out by 17 000 rcf centrifugation in 4°C for 2 min and resuspension of the pellet in PBS by brief water bath sonication for 3 times. The weight ratio of lipids for fusogenic liposome DOTAP:DODMA:DOPE:DSPE‐PEG (Avanti Polar Lipids; 890 890, 890 899, 850 725, and 880 120, respectively) was 22.25:22.25:44.5:11. The RNP concentration was 12 µM for in vitro application and 36 µM for in vivo application and comparison to LNP. The volume ratio of RNP:silica:lipid was 2:1:1.

4.3. Characterization of Nanoparticles

The prepared nanoparticles (0.1 mg mL⁻¹ about silica) were dispersed in PBS, and the concentration of liposome was adjusted to be same as concentration of lipid associated with silica in the lipid‐silica hybrid nanoparticle. Hydrodynamic diameter and zeta potential was measured using dynamic light scattering with a Malvern Zetasizer Nano ZS (Malvern, UK). The transmission electron microscopy samples were prepared by casting samples (10 µL) on a carbon film TEM grid (Ted Pella, 01840‐F) for 15 min, negative stained with UranyLess EM stain (Electron Microscopy Sciences, 22409) for 2 min and overnight drying at ambient condition. TEM images were acquired using JEM‐2100 transmission electron microscope (JEOL, Japan). The Brunauer‐Emmett‐Teller (BET) analysis of porous silica nanoparticles was conducted using ASAP2420 physisorption analyzer (Micromeritics Instruments, USA).

4.4. RNP Loading and Quantification

FAST‐CRISPR nanoparticles were prepared as above with different input RNP concentrations. Prepared nanoparticles were dispersed in RIPA buffer (Thermo Scientific, 89901) mixed with LDS sample buffer and reducing agents (Invitrogen, B0007 and B0009, respectively) at working concentration (50 µg mL⁻¹) and incubated for 10 min at 70°C. Samples were centrifuged at 17 000 rcf for 10 min and the supernatants were used for polyacrylamide gel electrophoresis at 200 V for 20 min using Bolt Bis‐Tris plus mini protein gel electrophoresis system (Invitrogen, USA). The polyacrylamide gel containing Cas9 protein was washed with deionized water, and Coomassie stained (Thermo Scientific, 24615) overnight and destained with deionized water. The stained gel was imaged using Bio‐Rad Gel Doc XR+ Gel Documentation system (Bio‐Rad, USA) and protein band intensity was analyzed using ImageJ software.

4.5. Cell Culture

The human colorectal cancer cell HCT‐116 (ATCC, CCL‐247) and embryonic kidney cell HEK293T (ATCC, CRL‐3216) were cultured in Roswell Park Memorial Institute 1640 Medium (RPMI, Gibco, 11875093) and Dulbecco's Modified Eagle Medium, high glucose (DMEM, Gibco, 11965092), respectively. All cell culture media were supplemented with heat‐inactivated fetal bovine serum (FBS, 10% v/v, Gibco, 10082147) and penicillin‐streptomycin (1% v/v, Gibco,151401222). Cell culture condition was maintained at 37°C in a humidified atmosphere with 5% CO₂ and subcultured using 0.25% trypsin‐EDTA (Gibco, 25200072).

4.6. Confocal Laser Scanning Microscopy Imaging of RNP Delivery

Cells were plated in 8‐well Nunc Lab‐Tek II Chambered Coverglass (Thermo Scientific, 155409) or Ibidi µ‐Slide (Ibidi, 80806) at a cell density of 30 000 cells per well in 300 µL media. The fluorescent gRNA was freshly prepared by annealing telomere‐targeting crRNA with ATTO 550‐labeled tracrRNA (Integrated DNA Technologies, 1075928) and used for RNP complex formation as above. The fluorescent RNP was delivered using FAST‐CRISPR nanoparticles or LNP, labeled with the lipophilic fluorescent dye DiD (Invitrogen, D7757) by adding to the lipid mixture in ethanol (1% w/w of total lipid) during formulation, at 2 days after cell plating and imaged with Hoechst 33342 (Invitrogen, H3570, 1:10 000) nucleus staining using Stellaris 5 (Leica, Germany) for incubation time dependent imaging and LSM980 (Carl Zeiss, Germany) for other applications after media change and analyzed using ImageJ software.

4.7. T7 Endonuclease I (T7E1) Assay

Cells were plated in 96 well‐plate (Corning, 3595) at a cell density of 8,000 cells per well in 100 µL media. The EMX1 gene editing RNP was delivered after overnight incubation, and cells were harvested 3 days later. The genomic DNA was extracted using DNA extraction solution (Lucigen, QE09050), the genomic locus containing the target site was PCR amplified using high‐fidelity PCR master mix (New England Biolabs, M0541) with primers listed in (Table S2; annealing at 68°C) and purified using PCR purification kit (Bioneer, K‐3627). The indel frequency was determined using T7 endonuclease I enzyme (T7E1, NEB, M0302) following the manufacturer's instructions. The DNA was analyzed by 2% agarose (Roche, 11685678001) gel electrophoresis, imaged using Bio‐Rad Gel Doc XR+ Gel Documentation system (Bio‐Rad, USA). The band intensity was measured using ImageJ software and indel frequency was calculated by

$$Indelfrequency\left(\%\right)=100\times \left(1-\sqrt{\frac{a}{a+b+c}}\right)$$

where a is band intensity of uncut DNA and b and c are intensities of cleaved DNA fragments.

4.8. Inhibition of Endocytosis

The HCT116 cell was prepared, and the fluorescent RNP was delivered using FAST‐CRISPR nanoparticle and CLSM imaged as above. The cells were pre‐incubated for 1 h with one of the endocytosis inhibitors, and media was replaced before RNP delivery: Genistein (GEN, 200 µM, inhibitor of caveolae‐mediated endocytosis, Sigma‐Aldrich, G6649), Chlorpromazine (CPZ, 20 µM, inhibitor of clathrin‐mediated endocytosis, Sigma‐Aldrich, C8138), Amiloride (AMI, 2 mM, inhibitor of macropinocytosis, Sigma‐Aldrich, A7410).

4.9. Immunofluorescence Staining of Cells

The HCT116 cell was plated on circular coverglasses (Paul Marienfeld, 0111500) in 24‐well plates (Corning, 3526) at a density of 50 000 cells per well in 500 µL media. The cover‐glasses were coated with poly‐D‐lysine (Gibco, A3890401, 1:10 in sterile water) by overnight incubation in a cell culture incubator and washed 3 times with sterile water before cell plating for cell adhesion. The DNA‐damaging RNP was delivered using FAST‐CRISPR along with the control groups after overnight culture of the cells. The cells were harvested 24 h later, washed once with PBS, and fixed with 4% paraformaldehyde (Biosesang, P2031) in PBS for 10 min at room temperature. They were washed with 0.2% Triton‐X 100 (Sigma‐Aldrich, T9284) in PBS (PBST) for 3 times for washing and permeabilization. The cells were incubated with 0.2% Triton‐X 100 blocking solution (Invitrogen, B10710) for 1 h at room temperature to prevent non‐specific antibody binding. The blocking solution was replaced with fresh 0.2% Triton‐X 100 blocking solution containing phospho‐histone H2A.X (Ser139, γH2AX) monoclonal antibody (3F2, Invitrogen, MA1‐2022, 1:500) and incubated overnight at 4°C. The cells were washed with PBST for 3 times and incubated with 0.2% Triton‐X 100 blocking agent containing Alexa 555 labeled goat anti‐mouse secondary antibody (Invitrogen, A‐21422, 1:500) for 2 h at room temperature. The cells were washed with PBS for 3 times, and nuclei were stained with Hoechst 33342 (Invitrogen, H3570, 1:10,000) and mounted on silane‐coated slide glass (Muto, 5116–20F) with mounting medium (Invitrogen, P36980) and imaged using CLSM.

4.10. Cell Viability Assay

Cells were plated in 96 well‐plate (Corning, 3595) at a cell density of 8000 cells per well in 100 µL media. The DNA‐damaging RNP was delivered with either FAST‐CRISPR or LNP, along with control groups, after overnight incubation, and cell viability assays were performed 3 days later. For CCK‐8 assay, media were replaced with fresh media containing 4% (v/v) of CCK‐8 reagent (Chromogen, CH‐3000) and incubated for 30 min and absorbance at 450 nm was measured using the Multiskan Sky microplate reader (Thermo Scientific, USA). The cell viability was calculated by

$$\mathrm{Viab}\mathrm{ility}\left(\%\right)=100\times \frac{A_{\mathrm{samp}\mathrm{le}}-A_{\mathrm{blank}}}{A_{\mathrm{cont}\mathrm{rol}}-A_{\mathrm{blank}}}$$

where A_blank is absorbance of media with CCK‐8 reagent without cell, A_control is absorbance of PBS control group and A_sample is absorbance of each experimental group. For LIVE/DEAD staining, cells were incubated with Hoechst 33342 for nucleus staining, followed by LIVE/DEAD staining reagent (Invitrogen, L3224, 1:2000 for calcein AM and 1:500 for ethidium homodimer‐1) for 30 min and imaged using Nikon ECLIPSE Ti2 microscope (Nikon, Japan).

4.11. Animal Experiments

All animal experiments were approved by the Institutional Animal Care and Use Committee at Ulsan National Institute of Science and Technology (UNISTIACUC‐24‐075). The xenograft tumor model was generated by subcutaneous injection of 5 × 10⁶ HCT116 cells suspended in PBS (100 µL) near the left flank of 5‐week‐old female BALB/C nude mice (Orient Bio, Republic of Korea). The tumor volume and body weight were monitored. The tumor volume was calculated by $V=\frac{a\times b^2}{2}$, where a and b are length of the long and short axis measured using calipers, respectively. When the tumor volume reached 100 mm³, mice were randomly divided into 3 groups of PBS, non‐target RNP@FAST‐CRISPR and target RNP@FAST‐CRISPR and particles (1 mg mL⁻¹ about silica) in PBS (60 µL) were intratumorally injected every single day for 12 days. Mice were euthanized with CO₂ the next day of last injection, and tumor tissues were harvested for further analysis.

4.12. Immunohistochemistry and TUNEL Assay of Tumor Tissues

The tumor tissue samples were fixed with 4% paraformaldehyde in PBS for 3 days at 4°C and cryoprotected in PBS containing 30% (w/v) sucrose (Sigma‐Aldrich, S9378) for 2 days at 4°C. The tissue samples were embedded in OCT compound (Sakura, 4583) and stored at −80°C and equilibrated to −20°C for overnight before sectioning. The samples were cryo‐sectioned in 10 µm thickness using a Leica cryostat (Leica, CM1950, Germany) and attached to silane‐coated slide glass and briefly dried at ambient conditions and stored at ‐80°C. For immunohistochemistry, the tissue sections were fixed with 4% paraformaldehyde for 10 min and washed with 0.4% Triton‐X 100 in PBS (PBST) for 3 times. Endogenous peroxidase activity was suppressed with 3% hydrogen peroxide (Sigma‐Aldrich, 216763) for 10 min and washed with PBST for 3 times. The tissue sections were incubated with 0.4% Triton‐X 100 blocking solution for 1 h at room temperature to prevent non‐specific antibody binding. The blocking solution was replaced with fresh 0.4% Triton‐X 100 blocking agent containing either phospho‐histone H2A.X (Ser139, γH2AX) monoclonal antibody (3F2, Invitrogen, MA1‐2022, 1:500) or Caspase 3/p17/p19 monoclonal antibody (2G4B2, Proteintech, 66470‐2‐Ig, 1:500) and incubated overnight at 4°C in a humidified chamber. The tissue sections were washed with PBST for 3 times and incubated with 0.4% Triton‐X 100 blocking agent containing HRP conjugated goat anti‐mouse secondary antibody (Invitrogen, 31430, 1:500) for 1 h at room temperature and washed with PBS for 3 times. The tissue sections were incubated with Pierce DAB substrate kit (Thermo Scientific, 34002) for 10 min and washed with PBS for 2 times and deionized water for 2 times. Tissue sections were washed for 15 min at each step. TUNEL assay was performed using colorimetric TUNEL assay kit (Abcam, ab206386) following manufacturer's instructions. The tissue sections were then incubated with Hematoxylin, Mayer's (Abcam, ab245880, 1:10 diluted in deionized water) for 2 min and washed with three changes of distilled water, and bluing reagent was applied for 15 s and washed with three changes of distilled water. The tissue sections were dehydrated in three changes of absolute ethanol and mounted in mounting medium (VectorLabs, H57000‐60) with coverglass, and images were acquired using BX51 optical microscope (Olympus, Japan) and analyzed using ImageJ software.

4.13. Data and Statistical Analysis

All data were represented as mean ± standard deviation, and statistical significance was analyzed using one or two‐way ANOVA or student's t test as indicated in each figure caption; ns, not significant; ^* p < 0.05; ^** p < 0.01; ^*** p < 0.001; ^**** p < 0.0001. GraphPad Prism software was used for analysis.

Conflicts of Interest

M.K. and J.J. have filed for intellectual properties based on the results of this study. E.J., K.M., T.K., S.W.C., and J.J. are shareholders of CasCure Therapeutics Inc. All other authors declare no potential conflicts of interest.

Supporting information

Supporting File: smll72145‐sup‐0001‐SuppMat.pdf

Data Availability Statement

All data that support this study are in the article and supplementary information available from the Wiley Online Library. Additional source data are available from the first and corresponding author upon reasonable request.