The pH-responsive nano-targeted protein degradation systems typically incorporate target protein degradation molecules into pH-sensitive nanocarriers. These systems remain stable under physiological conditions but undergo structural changes or chemical bond cleavage in acidic environments, enabling controlled release or activation. Building on this foundation, self-assembled pH-responsive nano-targeted protein degradation systems further integrate molecular self-assembly strategies with stimulus-responsive properties, representing a more refined and modular development direction in this field^[39] [Figure 2]. Importantly, beyond extracellular tumor acidity, intracellular pH gradients along the endocytic pathway (early endosomes pH ~6.0-6.5, late endosomes pH ~5.0-6.0, lysosomes pH ~4.5-5.0) provide additional opportunities for stepwise nanocarrier disassembly and cargo release. However, efficient endosomal escape remains one of the major bottlenecks limiting the therapeutic efficacy of pH-responsive nanocarriers^[40,41].

The pH-responsive targeted drug delivery mechanism utilizes the acidic microenvironment of tumors (extracellular pH ~6.5, intracellular pH ~4.5-5.5) as a core activation site. Drug targeting and release occur through three primary mechanisms^[42]. First, in polymers with high buffering capacity (e.g., polyamines), proton accumulation within endosomes induces the so-called “proton sponge effect”, in which continuous proton influx is accompanied by chloride ion entry and osmotic swelling. The resulting increase in osmotic pressure leads to endosomal membrane destabilization or rupture, thereby facilitating cytosolic release of degradation molecules^[43]. Second, acid-labile bonds (e.g., hydrazide, imine) between drugs and carriers or within carriers hydrolyze under acidic conditions, directly releasing drugs or facilitating intracellular release. Third, PEG detachment strategies responsive to pH utilize PEG coatings to maintain nanocarrier circulation under physiological conditions. In the acidic tumor environment, PEG detaches via acid-sensitive bond cleavage, resolving the “PEG dilemma” and improving intracellular uptake and drug release efficiency at the tumor site^[44,45].

The pH-responsive nano-PROTAC developed by the Xue group (MNC@Ca/MnCO₃/ARV/anti-PD1) gradually dissolves in the acidic TME (pH ≈ 5.5 - 6.5), releasing internally loaded PROTAC (ARV-825) and anti-PD1 molecules, which specifically degrades BRD4 without damaging immune cells. In neutral tissues (pH ≈ 7.0), the nanocarrier remains stable, minimizing drug release and reducing toxicity to normal cells^[46]. The universal nano-ATTEC platform based on mixed-shell polymeric micelles (MSPMs) designed by the Shi team overcomes poor cell-type specificity and dependence on intrinsic autophagy of conventional small-molecule ATTECs. The nano-ATTEC platform self-assembles from three diblock polymers: poly(ethylene glycol)-b-poly(ε-caprolactone) (PEG-b-PCL) modified with ligands for target proteins (e.g., BRD4 or AR), LC3 ligand (ispinesib, IS), and pH-responsive poly(β-amino ester)-b-poly(ε-caprolactone) (PAE-b-PCL). In acidic tumor environments (pH ≈ 6.5), protonation of the PAE segment promotes tumor cell uptake, autophagy activation, and LC3 targeting, leading to protein degradation (e.g., BRD4, AR) via the autophagy-lysosome pathway. Under normal physiological conditions (pH ≈ 7.4), the PEG shell shields the PAE segment, significantly reducing uptake by normal cells^[30].

Due to modular design, biocompatibility, and stimulus responsiveness, self-assembled nanocarriers have become ideal platforms for TPD systems (e.g., PROTACs, LYTACs, ATTECs). These nanocarriers self-assemble through non-covalent interactions (e.g., hydrophobic interactions, electrostatic interactions, π-π stacking), encapsulating or covalently conjugating degradation modules^[47]. Triggered by specific stimuli, they precisely activate or release degraders, enhancing targeting efficiency and reducing off-target toxicity. Tumors possess lower extracellular pH (~6.5) than normal tissues (~7.4), with even lower lysosomal pH (~4.5-5.0). pH-responsive nano-TPD systems exploit these differences by protonation or cleavage of acid-sensitive groups (e.g., hydrazone, acetal bonds, amino acrylates), releasing degradation modules specifically in tumor tissues^[48].

The bifunctional NLTC constructed by the Liu group addresses limitations such as insufficient tumor accumulation, non-specific activation, and low monotherapy efficacy. The lysosome-targeting bifunctional chimeric nanoplatform self-assembles into a nanocapsules structure with surface-modified lysosome-targeting ligands. Upon entering cells, protonation of amino groups in acidic lysosomes triggers structural dissociation, releasing LYTAC degraders. These degraders specifically degrade tumor cell surface proteins and activate antitumor immune responses. This system demonstrated a tumor inhibition rate of 87.9% in a 4T1 breast cancer model^[31]. To address the limitations of traditional LYTACs in hepatocellular carcinoma (HCC) therapy, such as low bioavailability and nonspecific distribution, the Chao group developed an intelligent modular DNA LYTAC (IMTAC) nanodevice and a drug EGFR/PDL1 IMTACs. Using self-linked DNA nanocircles as the carrier, this system achieves specific activation in the TME (pH 6.5-7.0) by precisely regulating the stoichiometry and modular distribution of pH-responsive degradation promoters (targeting insulin-like growth factor II receptor, IGFIIR) and EGFR/PDL1-targeting ligands, with an optimal spacing of 20 nm. The nanodevice enables synergistic degradation of EGFR and PDL1, reducing EGFR levels to approximately 10% in HepG2 cells in vitro and achieving a tumor inhibition rate of 80% in vivo, without significant toxic side effects^[49]. The Qi group has developed a novel ionizable nano-PROTAC (i16-ET^Nc) platform, which co-assembles a pH-responsive tertiary amine and amphiphilic conjugates. This combination enhances cellular uptake and endosomal escape via the proton sponge effect, overcoming traditional nano-PROTAC endosomal entrapment. The platform achieves potent degradation of leucine-rich α-2-glycoprotein 1 (LRG1) in cells (DC₅₀ = 17.1 μM and 19.5 μM for 4T1 and MDA-MB-231 cells, respectively). In vivo, it accumulates in tumors through EPR effect and active targeting, prolongs blood circulation (t_1/2 = 1.27 h), reduces tumor weight by 71.3%, and demonstrates no significant systemic toxicity, offering a universal strategy for TPD^[50].

This system’s advantages include its targeting capability and biocompatibility, it activates only under acidic conditions in diseased tissues or organelles, minimizing harm to normal tissues. Moreover, most acid-sensitive bonds comprise biodegradable materials, ensuring low toxicity of degradation products. However, substantial challenges remain, such as significant intratumoral pH gradient variations requiring precise control of acid-sensitive bond cleavage thresholds; endosomal escape efficiency affecting intracellular delivery and degradation by lysosomal enzymes; and inconsistent responses caused by pH fluctuations across patients or lesions.

The tumor microenvironment is frequently characterized by an aberrant protease landscape, in which multiple matrix metalloproteinases (MMPs) are upregulated in a wide range of tumor tissues and their surrounding extracellular matrix, and are closely associated with tumor invasion, metastasis, and disease progression^[51,52]. Enzyme-responsive nano-targeted protein degradation systems are typically designed by exploiting the abnormal expression and/or activity of tumor-associated enzymes - such as MMPs and cathepsins - within tumor tissues or along cellular endocytic pathways. In these systems, enzyme-cleavable substrates or linkers are incorporated into degrader prodrugs or nanocarrier architectures; upon encountering the target enzymes, site-specific cleavage occurs, thereby triggering the release and/or activation of degraders^[53].

The Ryu team developed a cancer-specific carbonic anhydrase IX (CAIX)-targeting supramolecular nanofibrous lysosome-targeting chimera (Supra-LYTAC). It comprises two hetero-functionalized, self-assembling amphiphilic Phe-Phe-Lys (FFK)-based peptides: one targets CAIX expressed on cancer cell membranes, and the other interacts with the POI. The supramolecular nanofibers near CAIX form a ternary complex (CAIX-nanofiber-POI). These complexes undergo clathrin-mediated endocytosis into lysosomes, where the acidic hydrolytic enzyme-rich environment facilitates the degradation of POIs^[54].

Tumor tissues overexpress specific proteases [e.g., MMPs, cathepsins (Cts), and plasmin]^[55]. Enzyme-responsive self-assembled nano-TPD uses protease-sensitive peptides as linkers or assembly units. Upon specific enzymatic cleavage at tumor sites, carriers undergo depolymerization or degradation module activation, enabling precise control.

The Gao team constructed dBET6@CFMPD nanomedicine by self-assembling MMP-2-sensitive peptide (FFRFKGPLGLAGC)-PEG-DSPE conjugates modified with the photosensitizer Ce6 (chlorin e6) and the BRD4-targeting PROTAC (dBET6). This system addresses multiple challenges: poor PROTAC solubility, insufficient tumor accumulation, limited single-agent therapy efficacy, immune escape caused by PD-L1 (programmed death-ligand 1) upregulation after photodynamic therapy (PDT), and impaired nanoparticle penetration and retention. At tumor sites, the highly expressed MMP-2 triggers structural transformation from spherical to nanofibrous morphology, enhancing Ce6 retention, PDT efficacy, and immunogenic cell death (ICD). Simultaneously, released dBET6 degrades BRD4, downregulates c-Myc and PD-L1 expression, reprograms tumor-associated macrophages (TAMs), and synergistically inhibits primary breast tumors and lung and brain metastases^[56].

The split-and-mix liposomal PROTAC system (LipoSM- PROTAC) self-assembles from folate (FA)-modified liposomes. Through DSPE-PEG2000, it links the E3 ubiquitin ligase ligand (VHL) and ERα-targeting ligand (4-hydroxytamoxifen, Tam), with MMP-2-sensitive peptides on the surface. Following MMP-2 cleavage at tumor sites, the liposome induces synergistic ligand activity, selectively degrading ERα protein and effectively inhibiting MCF-7 breast cancer cell growth^[57]. The targeted biomolecule regulation platform (SM-PROTAC) self-assembles into nanospheres with functional modules via diphenylglycine peptides (δδRR), encapsulating split PROTAC fragments (E3 ligand module and POI ligand module). After cellular uptake, the modules form active PROTACs in situ, targeting multiple proteins [e.g., ERα, cyclin-dependent kinases 4/6 (CDK4/6), bromodomain-containing proteins 2/4 (BRD2/4)] without significant activation in normal tissues^[58].

Enzyme-catalyzed processes exhibit amplification effects, as one enzyme molecule can cleave multiple carrier molecules, significantly enhancing sensitivity. Additionally, carriers enable complete control from stable circulation and tumor accumulation to stimulus-triggered release and degrader activation, reducing non-specific in vivo release and associated toxicity. However, enzyme activity may vary significantly among individuals, disease states, and tumor/inflammatory lesions. Thus, optimizing enzyme cleavage site sensitivity, carrier stability, and release kinetics based on actual enzyme expression remains critical. Furthermore, synthetic complexity, large-scale production, and quality control present barriers to clinical translation of enzyme-responsive nano-TPD. Although preliminary in vivo data demonstrate efficacy, systemic safety, localization and degrader release efficiency in complex microenvironments, and long-term effects still require comprehensive evaluation.

Under normal physiological conditions, intracellular redox reactions are typically maintained in a state of equilibrium^[59]. However, tumor tissues experience disturbances in oxidizing and reducing agents due to environmental factors and signal regulation, leading to redox imbalance. These disturbances subsequently influence tumor cell behavior, including metabolism, proliferation, metastasis, and drug resistance^[60].

ROS represent the main oxidizing agents inside cells. ROS are mainly generated by the mitochondrial respiratory chain during oxidative stress. Common ROS include hydroxyl radicals (•OH), superoxide anions (•O₂-), singlet oxygen (¹O₂), and H₂O₂^[61]. Under normal conditions, ROS remain at low levels (0.02-1 μM), supporting normal cell growth and apoptosis. However, when ROS levels significantly rise (50-100 μM) ,proteins associated with ROS-sensitive signaling pathways become activated, leading to oxidative stress in tumors or inflamed tissues^[62]. The main reducing agents within cells include GSH, Catalase (CAT), thioredoxin, and Fe²⁺. Among these, GSH has the highest cytoplasmic concentration in tumor cells (2-10 mM), greatly exceeding its extracellular concentration (2-20 μM)^[63]. GSH concentration in tumor cells is about 1000-fold higher than in plasma and at least four times greater than in normal tissues^[64].

Chen’s group developed ROS-responsive Pre-PROTACs, a type of ROS-responsive small-molecule conjugate, by introducing ROS-cleavable groups (aryl boronic acid) into the parent PROTAC molecules. These Pre-PROTACs efficiently degrade the target protein BRD3 in response to elevated ROS levels^[65]. Bi et al. designed a ROS-responsive nano-PROTAC (BT-NanoPROTAC) targeting BRD4, which is formed by the self-assembly of Small-Molecule Conjugate BT-PROTAC (ARV771 modified with a boronic acid group) and Poloxamer 188 into nanoparticles and remains stable in the bloodstream. This system leverages the high ROS levels in activated hepatic stellate cells (aHSCs) to release ARV771, a potent BRD4 degrader. The BT-NanoPROTAC specifically degrades BRD4 in aHSCs, downregulating fibrosis markers such as α-SMA and type I collagen, thus demonstrating therapeutic potential for hepatic fibrosis both in vitro and in vivo^[66]. The ROS-guided supramolecular assembly system (TCO-MZ1) contains a PROTAC precursor and ROS-sensitive monomers. In the ROS-rich tumor environment, oxidative reactions occur in these monomers, reorganizing micelles into more stable nanofiber structures and prolonging PROTAC retention in tumors. Concurrently, ROS triggers precursor activation, degrading PD-L1 and reversing the immunosuppressive microenvironment^[67].

The redox balance in tumor cells is disrupted, with GSH concentrations (10-20 mM) significantly higher than normal tissues (2-10 μM), accompanied by substantially elevated ROS levels. ROS/GSH-responsive self-assembled nano-TPDs employ sensitive linkers, such as disulfide and selenoether bonds, which break upon changes in the redox environment to release PROTACs. Glutathione-scavenging nanoparticles (ARV@PDSA Nano-PROTAC) possess a core-shell structure created by polymer self-assembly. The core encapsulates PROTAC agents, and the shell features a GSH-sensitive cross-linked disulfide network. Upon entering tumor cells, high intracellular GSH concentrations cleave the disulfide bonds, triggering carrier disassembly and PROTAC release. Simultaneously, these nanoparticles consume excess GSH, increasing oxidative stress and synergistically enhancing HER2 degradation. This approach achieves a 40% higher tumor growth inhibition rate compared to PROTAC treatment alone^[68]. To address the broad-spectrum cytotoxicity issues associated with molecular glues, the Yuan team designed a targeted molecular glue, H1-mGlu, against the Bcr-Abl protein, and constructed a self-assembling nanocarrier, Cle-NP. This system triggers drug release through high concentrations of GSH/H₂O₂ in the tumor microenvironment, achieving tumor-specific degradation. In vitro experiments demonstrated that Cle-NP exhibited a median inhibitory concentration (GI₅₀ = 74.8 nM) for K562 cells, effectively degraded Bcr-Abl (DC₅₀ = 0.89 μM), and induced apoptosis. In vivo, Cle-NP significantly inhibited tumor growth (TGI = 78.1%) in K562 xenograft models without causing notable liver damage or weight loss, resolving the systemic toxicity issues of traditional molecular glues and providing a novel strategy for precision cancer therapy^[29]. The self-assembled LYTAC nanosystem connective tissue growth factor (CTGF)-LYTAC (NanoCLY) is a GSH-responsive nanoparticle self-assembled from amphiphilic chimeras, which targets CTGF by forming nanoaggregates through peptide chain hydrophobic interactions, with tumor-targeting peptides on its surface. Upon entering the GSH-rich intracellular environment, the disulfide bonds linking the hydrophilic CL8-M6P and the hydrophobic lauryl group are cleaved, disrupting the amphiphilic balance of the chimeras and leading to the disassembly of the originally self-assembled spherical nanostructures, thereby releasing the functional drug molecule CL8-M6P. This process mediates CTGF degradation and remodels the inflammatory TME, significantly enhancing the therapeutic efficacy against triple-negative breast cancer^[69].

ROS-responsive nano-TPD systems are characterized by rapid response and high selectivity. However, ROS molecules have short half-lives in vivo (e.g., approximately milliseconds to seconds for H₂O₂). Additionally, variability in ROS types among different lesions (e.g., H₂O₂ predominant in tumors, superoxide anions predominant in inflammation) may impact carrier response efficiency and TPD^[70,71].

Overall, endogenous condition - initiated nano-TPD systems exploit intrinsic physicochemical and biochemical disparities between diseased and normal tissues to achieve passive targeting and stimulus-responsive activation without external inputs. Owing to their design simplicity and biological relevance, this activation mode remains the most extensively investigated strategy in nano-TPD research. However, its reliance on heterogeneous endogenous signals also highlights inherent limitations in controllability and activation precision, motivating the exploration of exogenously triggered and multi-condition - initiated nano-TPD systems discussed in the following sections.

Physical signal activation represents a pivotal strategy that triggers system responses through externally applied physical energy (e.g., light, ultrasound, or magnetic fields). Compared with endogenous microenvironmental response mechanisms, these systems are immune to individual variations or pathological heterogeneity, demonstrating excellent predictability and reproducibility^[77]. Furthermore, physical signal stimulation offers distinct advantages including non-invasiveness, remote control capability, and precise parameter adjustment. This enables researchers to initiate or enhance protein degradation processes at specific temporal and spatial locations, achieving high-precision modulation of target protein degradation. Such highly controllable stimulation methods provide novel opportunities to minimize systemic side effects and optimize therapeutic windows.

Chemically activated nanoscale protein degradation systems primarily utilize exogenous bioorthogonal reactions as their trigger mechanism. By introducing bioorthogonal chemical reagents or specific enzyme-substrate pairs, these systems induce highly targeted chemical reactions within complex biological systems to activate the degradation process. As bioorthogonal reactions do not interfere with endogenous biochemical processes, this strategy demonstrates distinct advantages including high reaction specificity and excellent biocompatibility, enabling precise and controlled protein degradation in cellular or in vivo environments^[105].

Currently, chemical signal activation strategies primarily fall into two categories: click chemistry-based approaches and enzyme-catalyzed orthogonal reactions. The former utilizes classical or modified click chemistry reactions to achieve in-situ assembly or conformational transformation of target molecules, while the latter relies on substrate conversion mediated by specific enzymes to enable spatiotemporal control of degradation activity^[106]. These strategies provide crucial tools for constructing highly selective nanoscale protein degradation systems in complex biological environments.

Dual-endogenous co-activation systems target two distinct pathological signals within the TME, including pH, enzyme activity, ROS, GSH, and copper ions^[121,122]. Leveraging synergistic effects of these signals provides dual verification, ensuring degradation initiation only occurs at the intended site. This strategy is particularly effective for malignancies characterized by clearly defined microenvironmental profiles, such as pancreatic and lung cancers^[123].

The endogenous - exogenous co-activated nano-TPD system integrates a nanocarrier core with dual activation modules: “endogenous TME signals” (e.g., pH, enzyme, hypoxia, or ROS/GSH) and “externally controlled exogenous signals” (e.g., light, ultrasound, or magnetic fields)^[128]. This technology triggers TPD via cascaded synergy between these two signal types. Its core advantage is “dual-signal-dependent precise activation”, exogenous signals provide controlled spatial and temporal initiation, while endogenous signals ensure specificity and amplify degradation effects within target regions. Requiring simultaneous activation by multiple signals improves targeting specificity, reducing off-target effects associated with single-signal fluctuations. By integrating endogenous targeting and exogenous control, the system achieves progressive “tumor localization → on-demand degradation”, adapting to complex pathological environments. This approach addresses TME heterogeneity by responding to multiple signals, covering broader pathological conditions. It also significantly reduces PROTAC exposure in healthy tissues, representing one of the most clinically promising systems currently available^[129].

Existing systems generally achieve synergy via two cascade modes: “exogenous signal localization → endogenous signal activation” or “exogenous signal triggering → endogenous signal amplification”. Relevant literature is classified by the type of exogenous signal applied.

Different nano-TPD systems vary substantially in core performance metrics, nano-TPD platforms should be benchmarked as a multi-objective trade-off^[139]. In endogenous initiation, activation is “device-free” and typically shows favorable biocompatibility, yet the activation threshold, kinetics, and effective dose are tightly coupled to the spatially heterogeneous TME, which can yield patient-to-patient variability and incomplete degradation in poorly perfused or “cold” lesions. This dependence also complicates cross-study comparability because the same formulation may behave differently across models with distinct microenvironmental distributions^[16].

In contrast, exogenous initiation decouples activation from TME variability by using externally applied signals that provide remote timing control and high spatiotemporal precision, thereby narrowing systemic exposure windows and potentially reducing off-target toxicity. However, these gains introduce new boundary conditions: tissue penetration depth and focal volume (notably for light), dependence on specialized equipment/clinical workflow, and the possibility of collateral activation when the stimulus field cannot be perfectly confined^[140-142].

Multi-condition initiation partially reconciles the above by implementing logic-gated degradation (AND/sequential gating): endogenous cues act as a biological “permission” signal, while an exogenous input provides temporal/spatial “authorization”. Such architectures can enlarge the safety margin for targets with narrow therapeutic windows by requiring multiple conditions before degradation occurs. Nevertheless, adding triggers and functional modules typically increases synthetic complexity, quality control burden, and batch-to-batch reproducibility risk, which may dominate translational feasibility^[136].

Finally, self-assembled nano-TPD systems reshape the performance profile by offering high apparent loading/modularity and improved intracellular delivery; yet their supramolecular heterogeneity, in vivo remodeling, and stimulus-independent leakage can confound mechanistic attribution. Therefore, future benchmarking should report not only tumor inhibition but also quantitative degradation kinetics (extent, onset, duration, reversibility) under clinically relevant conditions, alongside manufacturability and stability metrics^[143] [Table 3].

Although nano-TPD systems have made considerable progress in basic research, several significant challenges remain for clinical translation^[144]. These include balancing microenvironmental heterogeneity with individualized matching, ensuring equilibrium between carrier stability and response sensitivity, overcoming barriers in industrial-scale synthesis and quality control, and addressing issues related to pathway coordination and safety evaluation. Carrier fate (ADME), long-term toxicity, immunogenicity, and combination-product regulatory pathways should be explicitly addressed to strengthen the translational argument^[143-146].

A major challenge for endogenous activation systems is microenvironmental heterogeneity, which complicates personalized matching. Microenvironment parameters (e.g., pH, enzyme activity, ROS levels, GSH concentration) differ substantially among patients, disease stages, and within lesions themselves. Designing a universal response threshold based on “average characteristics” to ensure efficacy and safety across cases is difficult^[147]. Even in combined activation systems, precisely aligning patient-specific microenvironment features with exogenous stimuli remains challenging. Current methods for personalized assessment and customized design remain inadequate, highlighting the urgent need for precise, individualized treatment strategies to enhance therapeutic targeting and effectiveness^[148]. Notably, in clinically approved nanomedicines, improving circulation alone does not guarantee improved tissue accessibility; rather, human - animal mismatches in distribution/clearance and tissue “gatekeeping” frequently dominate outcomes^[149]. Therefore, nano-TPD designs that rely on microenvironment triggers should be coupled with biomarker-based stratification and Pharmacokinetics/Pharmacodynamics (PK/PD)-linked activation modeling to avoid under-activation (loss of efficacy) and over-activation (off-target toxicity) in heterogeneous clinical populations.

Balancing carrier stability with response sensitivity also represents a critical research challenge. Nano-TPD systems must maintain stability during systemic circulation yet rapidly respond at target sites. In self-assembled systems, factors such as plasma protein adsorption and blood shear forces can destabilize carriers, causing premature drug release or structural changes. In exogenously responsive systems, enhanced sensitivity can improve therapeutic outcomes but may compromise circulation stability^[150-152]. Conversely, prioritizing stability may reduce response speed. Thus, identifying an optimal balance between these two aspects remains essential^[153]. This balance should be evaluated within a full ADME framework - where the key questions are “Where does the carrier go? How long does it persist? How does it transform in vivo? How is it cleared?” A Nature Reviews Bioengineering review emphasizes that controlling blood circulation, biodistribution, tissue accessibility, and clearance is central to nanomedicine translation^[149]. For nano-TPD, such ADME uncertainty directly affects degrader exposure, intracellular delivery, and ultimately degradation kinetics, making it insufficient to report only in vitro responsiveness or short-term tumor accumulation.

Safety concerns, such as long-term residues from inorganic nanocarriers and immunogenicity of orthogonal enzymes, lack comprehensive long-term toxicology data, undermining assurances of biosafety^[154-156]. Clinically approved liposomal products (e.g., Doxil) demonstrate that nanomedicines can be successful, but also highlight that release testing, stability, and bioequivalence-sensitive quality attributes can be rate-limiting for regulatory review and lifecycle management^[157]. FDA has specifically developed science and guidance around nanomaterial-containing products and liposome drug products, reflecting the regulatory focus on Chemistry, Manufacturing, and Controls (CMC), human PK/Bioequivalence (BE), labeling, and in vitro release testing for complex nanosystems^[158]. Immunologically, PEGylated nanocarriers can trigger anti-PEG antibodies, leading to the accelerated blood clearance (ABC) phenomenon upon repeat dosing, reducing efficacy and altering safety^[159]. In addition, liposomes and other nanoparticles may cause complement activation-related pseudoallergy (CARPA), an unpredictable hypersensitivity risk that has been discussed for decades and remains clinically relevant for reactogenic nanocarriers^[160,161]. These well-documented phenomena imply that nano-TPD platforms - especially those intended for repeated administration - should incorporate immunotoxicity endpoints (complement activation, cytokine release, anti-carrier antibodies, ABC risk) early in development rather than treating them as late-stage toxicology add-ons^[162].

For TPD itself, safety is not only material-driven: off-target degradation and system-level consequences of degrading proteins require proteomics-enabled selectivity/safety profiling to reduce unintended toxicity^[163]. Thus, long-term biosafety packages should integrate both carrier-centric (retention/accumulation, chronic inflammation/fibrosis) and mechanism-centric (degradome/off-target) assessments.

Industrialization barriers in synthesis and quality control significantly hinder the clinical application of nano-TPD systems. Self-assembled and co-activated systems typically integrate multiple functional components (e.g., sensitive linkers, targeting ligands, responsive modules, TPD loading sites), resulting in complex structures, lengthy synthesis routes, difficulty in intermediate purification, and poor reproducibility between batches. During large-scale production, these complexities not only increase synthesis difficulties but also complicate establishing reliable quality control standards^[164]. Furthermore, accurately characterizing nano-TPD behavior in vivo requires simultaneous monitoring of multiple processes, including carrier distribution, structural changes, TPD release, and target protein degradation. Current characterization methods lack real-time, in situ evaluation capabilities, limiting clinical applicability^[165]. Regulators expect sponsors to define critical quality attributes (CQAs) and link them to performance. FDA guidance on drug products containing nanomaterials outlines development expectations, while liposome-specific guidance highlights CMC and human PK considerations; FDA’s continued work on nano-size complex product in vitro release testing underscores how release/transformations become pivotal for complex nanoformulations. Therefore, nano-TPD manufacturing should move toward QbD-driven process control, orthogonal characterization (size distribution, morphology, surface chemistry/ligand density, encapsulation vs. free drug, stability, release/activation kinetics), and clinically relevant comparability plans^[166].

Finally, coordinating treatment pathways and ensuring safety in clinical translation are critical. Exogenous and co-activated systems typically require specialized equipment and professional operation, complicating therapy coordination^[167,168]. Dosing schedules, pre-activation windows, and external stimulus parameters increase operational complexity and patient compliance challenges^[167]. Regulatory evaluation criteria for “nanomedicine + external device” co-activation models remain incomplete, delaying clinical translation. FDA defines “combination products” (drug/device, biologic/device, etc.) and provides cGMP and premarket pathway principles for such products, underscoring that integrated drug - device development must address not only pharmacology but also device controls, quality systems, and review coordination^[169]. Hence, nano-TPD co-activation platforms should pre-emptively plan for combination-product requirements: standardized stimulus parameters, operator-independent workflows, device validation, and integrated risk management.

Overall, advancing nano-TPD systems clinically requires technological breakthroughs, rigorous safety assessments, scalable manufacturing with defined CQAs, and standardized clinical/regulatory protocols to ensure feasibility and safety^[170,171].